Physical evaluation of computed radiography as a ... - BIR Publications

1994, The British Journal of Radiology, 67, 988-996

Physical evaluation of computed radiography as a mammographic X-ray imaging system A W O R K M A N BSc, MSc, A R COWEN, BSc and D S BRETTLE, MSc FAXIL, University of Leeds, Academic Unit of Medical Physics, Leeds General Infirmary, Leeds LS1 3EX, UK

Abstract The physical imaging properties of a computed radiography (CR) system operating under mammographic exposure conditions have been measured. These measurements include modulation transfer function (MTF), sensitometric response and noise power spectrum (NPS). These figures were used to derive signal-to-noise ratio (SNR) descriptors of the system performance. The same measures were derived for a mammographic film-screen system. The SNR properties of the CR system indicate superior low frequency performance over a wider dynamic range than the film-screen system. The spatial frequency dependent SNR properties of the CR system are, however, limited. The results of a psychophysical test support the results obtained from physical measurements. These results indicate that despite its limited sharpness, details of dimensions relevant to diagnosis in mammography (circular details 0.25 mm diameter to represent microcalcifications and 3 mm diameter to represent small masses) are reproduced using CR with SNRs which are at least comparable with those of a mammographic film-screen system.

Mammography is an accepted procedure for the early detection of breast cancer. When applied to a large population of asymptomatic patients it has the potential to reduce the mortality rate by a third or more [1]. The earliest mammographic indicators associated with breast disease are often very subtle features such as low contrast small area mass lesions, micro-calcifications and architectural distortions. Additionally, the background morphology of the breast can vary from low contrast uniform areas to very high contrast busy areas. These factors make mammography one of the most demanding radiological examination both diagnostically and technically. In screening mammography the radiologist is required to read large numbers of mammograms derived from an asymptomatic population. If such a programme is to be effective a sufficiently high proportion of patients presenting with pathology must be detected. This requires not only sufficiently good image quality to be able to detect pathological indicators but also reliable and consistent interpretation of the mammograms. Digital mammography has for a long time been identified as having the potential to alleviate the demands of mammography [2]. The potential benefits include: computer aided diagnosis, teleradiology and digital archival/ communications. The clinical application of digital mammography beyond the research level has to date Received 11 October 1993 and in final form 11 March 1994, accepted 28 March 1994. 988

been limited. This can be particularly attributed to two main factors; firstly, the lack of a suitable digital acquisition system capable of producing images of a clinically acceptable standard and secondly, the high demands on computer resources required to manage, store, process and display images of a suitable specification. Indeed the pixel size required to maintain the diagnostic information in digital mammograms has been and still is the subject of some debate [3]. The acquisition of digital data has historically been through the digitization of conventional film-screen images, usually by TV camera, laser film scanner or film densitometer. There are, however, two major problems inherent in film digitization; firstly, the digital image is restricted by the physical limitations of the film itself particularly with respect to dynamic range and secondly, the digitization stage adds to the image data its own sources of image degradation [4]. With the increased availability of computed radiography (CR) [5] over the last few years interest in digital mammography utilizing direct digital acquisition has grown. Furthermore, improvements in image quality with successive generations of system hardware and storage phosphor plate technology [6, 7] have encouraged more workers to investigate the potential of CR as an image acquisition system for digital mammography. We have been involved for the last 3 years in an evaluation of digital mammography using CR as the image acquisition system. This evaluation has included test object [8] and clinical evaluations [9] which indicate that CR produces diagnostically comparable results to The British Journal of Radiology, October, 1994

Evaluation of CR as a mammographic X-ray imaging system

the system, with a latitude of 2.0 defining a range of 100 {i.e. log10 of the exposure range). The sensitivity is a measure of the gain of the reader system and is analogous to a speed measure. The latitude wasfixedat 2.0 and the sensitivity fixed at 200. This allowed the image pixel value to change with varying exposure to the plate. Screen transmission measurements were made using a PTW Dali exposure meter and PTW 77334 1 cc ionization chamber which has a flat response at mammographic beam energies. Measurements were made utilizing the same beam quality as had been used for the X-ray sensitometry. The screens (removed from the cassettes) Method and materials A Philips Graphic He Computed Radiography system were placed close to the tube exit port and the ionization installed in the Diagnostic X-ray Department of Leeds chamber placed at a distance of 1 m from the tube. General Infirmary was employed in our study. This Films for noise power analysis were exposed in a system incorporates the Fuji Computed Radiography manner similar to sensitometric exposures except that to 7000 series image reader. In our study third generation obtain sufficient area of uniform exposure the beam high resolution image plates HR III and the more recent profile was flattened by using a PMMA wedge [11]. The generation HR IIIN plates were employed. These plates variation in PMMA attenuation over the area of interest have nominal dimensions of 18 x 24 cm and the (12 x 12 cm) produces only a small change in the specsampling matrix used is such that the digital sampling trum incident on the receptor (measurements show a interval is 0.1 mm. This produces a digital image with change of less than 2% in the number of quanta yuGy"1). dimensions of 1770 x 2370 pixels. The signal output is Image plates for noise analysis were exposed in a similar digitized to 10 bits of grey scale following logarithmic manner to the film-screen systems and read 3 min amplification. For the purpose of analysis digital image following exposure. Plates were exposed to six different data were transferred over a fibre optic ethernet link to levels encompassing approximately a 100 fold range in a Sun SPARCserver 630 MP located in the FAXIL receptor entrance exposure. Image plates were read in a laboratory. SEMI-AUTO reading mode which had fixed latitude We have also undertaken an image quality study of (2.0) and variable sensitivity [12]. The noise power the mammographic film-screen system routinely used in spectrum of the CR images were measured by a method Leeds General Infirmary. Thefilm-screenused was Fuji previously outlined [13]. The noise power spectrum of HR Mammo Fine with CEA MA single emulsion film. the mammographic screen-film images were measured Thisfilm-screensystem was used as the gold standard in by a method previously used to measure the NPS of the clinical study for comparison with CR, with medio- conventional film-screen systems [12]. Uncertainties in lateral and craniocaudal views being acquired on film- the NPS measurement were typically 5% and were screen and the mediolateral oblique view on CR. The mainly determined by the number of noise traces used in mammographic films were processed in a Kodak M8 the NPS determination. processor using a standard 90 s processing cycle using The pre-sampling MTF of the HR III and HR IIIN Rapitech developer and fixer at 34°C. It is recognized plates were measured in directions parallel and perpenthat the processing regime used in our study and the dicular to the laser scan direction using an angulated clinical study is different to that which may be used at slit technique [14]. The (MTF) of the mammographic dedicated mammography facilities and that these con- film-screen system was measured using a specially ditions may therefore not be the most optimal. Expo- constructed resolution test grating [15] with square sures were performed on a Philips MammoDiagnost UC wave frequencies ranging from 0.05 lp mm"1 up to dedicated mammographic X-ray unit. This incorporates 201pmm"1. The modulation of the bars and spaces at a Mo target X-ray tube with 30 /im Mo filtration. The the various spatial frequencies were measured using a half value layer of the beam with the compression plate scanning microdensitometer (Joyce Loebl type 3CS). in place was measured at 0.36 mmAl. The generator kV Results were corrected using the sensitometric curve of calibration was checked using an RTI PMX-III non- the film and converted to sine wave response. The invasive multimeter. All test exposures were made with modulations were then normalized to that of the the antiscatter grid removed. X-ray sensitometric ex- 0.05 lp mm ' group. Typical uncertainties in the MTF posures were performed at 28 kVp with the beam filtered determination estimated from the reproducibility of to simulate the filtration due to a compressed breast of results were 3%. average thickness. This was simulated using 4 cm of A two alternative forced choice (2-AFC) test [16] was PMMA [10] which was placed on the tube port. This used to determine the detectability of details of diagnosbeam quality corresponds to a X-ray quantumfluenceof tically relevant dimensions for each of the imaging 5 900 quanta mm"2 for an air kerma of 1 jiGy. CR image systems. The probability of a correct response in this test plates were read in a FIX mode where the CR reader is equal to the area under the receiver operator characparameters latitude and sensitivity are fixed quantities. teristic. All modalities tested were given the same recepThe latitude defines the range of exposures imaged by tor entrance air kerma (82 ^Gy). Exposure conditions film-screen systems for mammography. This paper presents the results of our investigation into the physical imaging characteristics of CR at mammographic beam qualities. We report measurements of the modulation transfer function (MTF) and noise power spectrum (NPS) of the storage phosphor system and derive from them the signal-to-noise ratio (SNR) measures, noise equivalent quanta (NEQ) and detective quantum efficiency (DQE). These measurements are also reported for a mammographic film-screen system.

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A Workman, A R Cowen and D S Brettle

were the same as those used for noisefilms/images.The CR images were processed with the processing algorithm which is applied to clinical mammographic images before printing onto laser hardcopy film [8]. The mean optical density for each film was 1.65. Pairs of images were obtained, one image in each pair containing a detail, the other noise only. The dimensions of the image regions were 4 x 4 cm. In the signal images the detail was placed at the centre of the image. Details of 3 mm [17] and 0.25 mm diameter were used in this study. The 0.25 mm detail was chosen to be representative of microcalcifications. Olson et al [18] analysed 52 clusters from 48 mammograms and reported 53% of malignant calcifications to be up to 210 ^m in size and 28% between 210 /xm and 420 /zm. The 0.25 mm diameter detail whilst not being at the resolution threshold of calcification sizes is representative of early stage disease. For each detail size, two different X-ray contrasts were used. 70 image pairs were obtained for each detector system for each detail size and contrast. Pairs of images (noise and signal + noise) were presented to observers who were required to indicate which image region contained the signal. Images were viewed on a light box masked to the test area. Observers underwent a training session before each viewing session where they were instructed on the size, contrast and possible position of the detail on each film. Observers were allowed free choice of viewing distance and were provided with a magnifying glass, however viewing time was limited to 30 s per film. Six observers viewed each of the image pairs. The standard errors in the calculated mean fractions of correct responses was determined by the method of Swets and Pickett [19]. Results and discussion

The characteristic curve of the Fuji HR Mammo Fine/CEA MA system is shown in Figure 1. This system requires a receptor entrance air kerma of 58/xGy to produce a density of one above base + fog level and gives a maximum gamma (y) of 3.49. The range of exposures this system can handle is limited by the properties of the radiographic film. Visibility at low exposures is limited in the toe region of the curve by lack of density and reduced film gamma. Visibility in the high exposure regions is limited by high density (although hot lights are used in mammography) and reduced contrast transfer in the knee region of the curve. Furthermore apart from any imaging constraints involved with the latitude/contrast trade-off properties of the film, filmscreen detectors also depend on the proper functioning of the automatic exposure system. Clinical CR images are usually acquired under an AUTO reading mode where the diagnostically relevant signal range is selected from the wide exposure latitude of the CR system (~ 104) [12]. This process involves a pre-reading mechanism which reads the image with the full latitude available in CR. From these data the range of intensities which compose the image are selected and the image plate undergoes the main reading process usually with a much reduced latitude (~ 102). The grey-scale reproduction of 990

3.5-

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Film Sensitometric Curve CR Image Processing Curve

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Figure 1. Sensitometric curves for Fuji HR mammo fine/CEA MA ( ) and computed radiography mammographic processing ( ).

an image in CR over a wide exposure range does not depend on reliable auto-exposure functioning but depends on the ability of the automatic read mechanism to detect the relevant information within the pre-read image data. From our clinical studies the usual distribution of latitude values for CR mammograms range from 1.6-2.3 with most being in the range 1.9-2.0. The relationship between the CR reader sensitivity value (S) and exposure to the image plate has previously been examined for normal diagnostic beam qualities [20]. The relationship between sensitivity and receptor exposure for mammographic beam qualities is shown in Figure 2 for the HR III and HR IIIN plates. These results were obtained from the exposures for noise power analysis. These indicate that both systems can acquire images with 1,000 HR-III Image Plate HR-IIIN Image Plate

0.01

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Air Kerma (mGy) Figure 2. CR image reader sensitivity value as a function of exposure for HRIII ( ) and HRIIIN ( ) image plates. The British Journal of Radiology, October, 1994


normal latitudes with average receptor entrance air kermas in the range 10/xGy to 1 mGy. For the same receptor exposure the HR III system gives a higher sensitivity value than the HR IIIN system. This indicates that for equivalent exposure levels the HR IIIN plates gives a higher stimulated light output than the HR III plates. CR data are usually processed prior to display presentation either on a viewing monitor or on hardcopy film. In this process the image data is passed through one of a selection of non-linear look-up tables which map pixel values tofilmdensity or monitor luminance. This results in an image with a grey scale rendition similar to that which would be obtained with a film-screen image. These curves can be further manipulated to change contrast, latitude and overall density to give any desired appearance. The shape of the standard curve used to process mammographic images is also shown in Figure 1. This presentation for clinical CR mammograms had been optimized using a test object in a previous study [8]. The X-ray absorption of the different system estimated using transmitted exposure measurements with an ionization chamber are shown in Table I. The CR image plates show higher absorption than the conventional radiographic screen. This may be attributed to both the different composition of the screens (CR image plate-barium flurohalide and HR mammo fine-gadolinium oxysulphide) and differences in the screen coating weights. The MTFs of the CR image plates and the film-screen system is shown in Figure 3 and indicates the superior signal transfer bandwidth of thefilm-screensystem. This results from the combination of a thin single screen and single emulsion film. However, as with any film-screen system, this is achieved by compromising the quantum detection efficiency of the detector. The MTF of the CR system does not depend on system speed in the conventional sense as images with the same sharpness may be produced at a wide range of exposures. Film-screen systems usually show reduced MTF as speed increases. The MTF of the CR images may be further manipulated by spatial frequency processing to increase the apparent sharpness of the CR images. The degree to which improvement may be achieved will obviously be limited by the signal-to-noise ratio of the system. The results also show that there is no significant difference in MTF between HR III and HR IIIN image plates. The spatial resolution limit of the CR images is limited by the sampling rate to 5 line pairs mm"1. The spatial resolution limit of thefilm-screensystem was measured at 20 line pairs mm"1. Table I. Radiographic screen absorptions measured at 28kVpMo spectrum filtered with 30/imMo and 4 cm PMMA Radiographic screen

Absorption

Fuji HR mammo fine CR HR III image plate CR HR IIIN image plate

0.76 0.83 0.86

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Fuji HR Mammo Fine/CEA MA HRIIIN HRIII

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Spatial Frequency [c/mm] Figure 3. Modulation transfer function for Fuji HR mammo fine/CEA MA and computed radiography HR III and HR IIIN images plates.

The NPS of a photostimulable phosphor system can be expressed as the sum of the NPS from X-ray quantum dependent sources, luminescent photon noise, screen noise and reader electronic noise [6, 7]. The NPS of a film-screen system can be expressed as the sum of the NPS from quantum dependent sources, film granularity noise and screen noise [21-23]. The noise power spectral values for the HR III and HR IIIN systems at 0.47 cycles mm"1 and 4.0 cycles mm"1 as a function of exposure to the image plates is shown in Figure 4. With increasing exposure to the image plates the noise power of the 1E-4 HR-III Image Plate HR-IIIN Image Plate

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Figure 4. Computed radiography H R I I I ( ) and H R I I I N ( ) noise power density versus incident air kerma for 0.47 cycles mm" 1 and 4 cycles mm" 1 spatial frequencies.

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systems decreases. In the lower exposure region this decrease is proportional to the increase in exposure indicating the limiting noise source arises from fluctuations in the number of X-ray quanta absorbed per unit area by the image plate. At higher exposure levels the noise power decreases less rapidly and tends toward a constant level. This level defines the noise power of residual noise sources arising from the fixed pattern granularity of the image plates and reader electronic noise sources. Measurements indicate that the HRIIIN system exhibits approximately half the residual noise level of the HR III system. It has been identified that the major contribution to the residual noise arises from the fixed pattern granularity of the image plates [7]. In this study the reader noise was identical for both plate systems, so that the reduced residual noise of the HR IIIN system must arise from reduced granularity of the HR IIIN image plate. The quantum noise level of all phosphor screen based imaging systems is enhanced beyond that expected based on the number of X-ray quanta absorbed per unit area by the system. This excess noise is due to variations in the amplification processes associated with the conversion of absorbed X-ray energy to light photons. Changing screen optical parameters such as introducing reflective backings and adding dyes to increase image sharpness will affect the excess noise level [23]. Excess noise in the storage phosphor system arises from both variations in the conversion process of X-ray energy to stored electrons and also depth dependent gain variations in the image readout process [24]. The noise power due to luminescent photon noise has been shown to make a significant contribution to the overall noise power of photostimulable phosphor systems [6, 7]. This is particularly so for the high frequency region of the NPS as the luminescence noise has essentially a white spectral distribution whilst the power of other noise sources decreases with increasing spatial frequency. The noise effective gain (a measure of the effective quantum gain of the system which would give the same level of secondary quantum noise as that measured) of the CR system for the HR III and HR IIIN image plates may be estimated by assuming that the high frequency noise power is due to the luminescence noise alone [6]. This givesfiguresfor the noise effective gain of 3.74 and 4.74 for the HR III and HR IIIN image plates, respectively. The higher noise effective gain of the HR IIIN system is due to the increased luminescent light output of the HR IIIN plate compared with the HR III image plate. This increase in light output can be estimated from the ratio of system sensitivity values. This indicates an estimated 25% increase in light output from the HRIIIN system. This figure agrees well with the 29% increase estimated from the luminescent photon noise levels. The noise equivalent quanta may be determined from knowledge of noise power spectrum, MTF, and macroscopic signal transfer gradient, y, and is a function of both exposure level and spatial frequency. 992

(ylo gl0 e) 2 -MTF(f) 2 NPS(f) NEQ is equivalent to the square of the signal-to-noise ratio of the acquired image data and may be related to the image quality since improved image SNR results in improved detectability given adequate display conditions for the observer to perceive the improvement. NEQ as a function of exposure and spatial frequency is mapped out for the three imaging systems in Figure 5. The contours map-out lines of equal noise equivalent quanta. The NEQ of the HR III and HR IIIN systems increase as a function of exposure even up to a receptor entrance air kerma of 1 mGy. For equivalent exposures the HR IIIN system provides higher NEQ values than the HR III system. The NEQ of the film-screen system increases as a function of exposure to a maximum and then reduces as exposure increases further. A more useful measure for comparing the performance of different imaging systems under identical exposure conditions is the detective quantum efficiency. This is a measure of how efficiently the imaging system uses the incident X-ray quanta. As such it is a measure of the efficiency of SNR transfer by the system since the SNR2 of the beam incident on the detector is given by the number of incident quanta per unit area (Q ) and the SNR2 in the acquired image is given by NEQ then: NEQ(f) =

NEQ(f) Q DQE as a function of exposure and spatial frequency is shown for the three imaging systems in Figure 6. The contour lines map-out lines of equal detective quantum efficiency. The maximum low frequency DQE of the film-screen system is lower than that of the CR image plates. However, the spatial frequency dependence of the DQE of the film-screen system does not fall off as rapidly as that of the CR system. Furthermore, it can be seen that the DQE of the CR system maintains its maximum value over a much wider exposure range than that of the film-screen system. The film-screen system maintains a high DQE level over a rather narrow exposure range. This is because the signal and noise properties of the intensifying screen are modulated by the contrast transfer function (y) of the film which is a function of the exposure level and the film grain noise increases with exposure (density). Therefore the use of a hot light to extend the dynamic range of the film may present areas of high density at a more comfortable viewing luminance; however, the efficiency of the recorded image is still compromised. The DQE of the CR system falls off at higher exposure levels due to the increasing dominance of image plate structure noise and residual CR reader system based noise components [7]. The low frequency DQE of film-screen systems is also partly compromised by increased film noise at low spatial frequencies, a factor which has been recognized by other workers [21,22]. The high quantum fluence necessary in mammography results in reduced quantum noise leading to the increased relative importance of film DQE(f) =

The British Journal of Radiology, October, 1994


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Figure 5. Noise equivalent quanta as a function of spatial frequency and air kerma for Fuji HR mammo fine/CEA MA and (a) HR III, (b) HR IIIN.

noise which not only limits low frequency DQE but also severely limits high frequency DQE due to the frequency independent nature of the film noise power at higher spatial frequencies. It is seen that the maximum DQE of the film-screen system occurs at exposures lower than that required to produce a net density of one. This has also been observed for several other mammographic film-screen combinations [25]. The results of the 2-AFC experiments are shown in Figure 7. These show the mean value of the fraction of correct responses for the 3 mm and 0.25 mm diameter details plotted against signal contrast for the three imaging systems evaluated. The error bars represent one standard deviation of the estimate. Results for only one contrast are shown for the 3 mm details as the higher contrast detail produced detectabilities close to 100% for all three systems and the results were therefore discarded. Although the results for the three systems do not indicate a statistically significant difference at the 5% level, a number of general observations can be made. Results indicate that for the 3 mm details the highest detectability is recorded for the HRIIIN system followed by the HR III system with the film-screen system giving the lowest detectability. Results for the 0.25 mm details show that the HRIIIN system exhibits the highest detectability followed by that of the film-screen system with the HRIII system showing the lowest Vol. 67, No. 802

detectability. For circular details of diameter d frequencies n e a r / « 1/2J contribute mainly to detection [26]. The performance of the three detector systems can be reconciled with their frequency dependent DQE behaviour at the exposure level used to acquire the images. Detectability will, however, also depend on the details being presented at sufficient contrast to allow the observer to fully utilize the SNR of the acquired image. Although the CR mammogram printed on film is processed with a look-up table to present an image which resembles a film-screen mammogram presentation, the CR image may be re-processed at will and this effectively allows the viewer access to the full available NEQ of the acquired image over its dynamic range and therefore allows optimization of signal detectability. However, this may not be the case with the film-screen mammogram as information is presented with an exposure level dependent contrast transfer factor and under some circumstances it is possible that signals may not be detected due to lack of contrast and not lack of SNR. Conclusions

The necessary specification for a digitized mammogram has been investigated by other workers [27,28]. They indicate that a 2048 x 2048 image with an effective pixel dimension of 100 ^m and 1024 grey levels to be an acceptable minimum specification. This matches the 993


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(a) Figure 6. Detective quantum efficiency as a function of spatial frequency and air kerma for Fuji HR mammo fine/CEA MA and (a) HR III, (b) HR IIIN.


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Figure 7. Detectability of details for Fuji HR mammo fine/CEA MA, HR III and HR IIIN for (a) 0.25 mm diameter and (b) 3 mm diameter details.

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Evaluation of CR as a mammographic X-ray imaging system specifications for the resultant digital image for CR mammography. Critics of CR quote the low spatial resolution limit as being deficient for the detection of early micro-calcifications. However, our experience and that of others [9,29-31] indicate that although the rendition of calcifications is lower for CR the detection rate is equal for both the modalities. Additionally, it should be noted that micro-calcification portrayal should not be the only or even the most important imaging goal of a digital mammographic system. Rather, since the more rapidly growing breast cancers present mammographically as poorly defined masses, it will be especially important for digital mammography to detect small masses in order to have the greatest impact (Sickles, E, Private Communication, 1992). The advent of high powered low cost computers and high resolution displays coupled with CR as an acquisition device has brought digital mammography into the clinical situation. The results presented in this paper illustrate the physical properties of CR that enable it to produce comparable results to film-screen systems, even though it has arguably deficient spatial resolution for mammography. Successive generations of CR image plate technology have shown improved per-formance in the mammographic imaging field. Future generations of CR plate (HR V) may provide further improvements. Also new photostimulable phosphor materials are under investigation which promise even better imaging performance for computed radiography in mammography [32].

Acknowledgments The authors would like to acknowledge the support of the United Kingdom Department of Health Medical Devices Directorate and the radiological staff of the Diagnostic Radiology Department of Leeds General Infirmary.

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7. WORKMAN, A and COWEN, A R, Signal, noise and SNR transfer properties of computed radiography, Phys. Med. Biol., 38, 1789-1808 (1993). 8. COWEN, A R, BRETTLE, D S, PARKIN, G J S and COLEMAN, N J, A preliminary investigation into the imaging performance of photostimulable phosphor computed radiography using a new design of mammographic quality control test object, Br. J. Radiol., 65, 528-535 (1992). 9. BRETTLE, D S, WARD, S C, PARKIN, G J S and COWEN, A R, A clinical comparison between conventional and digital mammography utilizing computed radiography, Br. J. Radiol., 67, 464^68 (1994). 10. IPSM, Report 59, Commissioning and Routine Testing of Mammographic X-ray Systems (IPSM, York) (1989). 11. COWEN, A R, BRETTLE, D S and WORKMAN, A, Compensation for field non-uniformity on a mammographic X-ray unit, Br. J. Radiol., 66, 150-154 (1993). 12. FAXIL, Report: A Physical Evaluation of Computed Radiography (MDD/92/36) (Department of Health, London) (1992). 13. COWEN, A R and WORKMAN A, A physical image quality evaluation of a digital spot fluorography system, Phys. Med. Biol., 37, 325-342 (1992). 14. FUJITA, H, MORISHITA, J, UEDA, K ET AL, Resolution properties of a computed radiographic system, Proc. SPIE, 1090, 263-275 (1989). 15. FAXIL, Report: The Imaging Sharpness of Mammographic Screen Film Systems (STD/89/17) (Department of Health, London) (1990). 16. GREEN, D M and SWETS, J A, Signal Detection Theory and Psychophysics (Peninsula Publishing, Los Altos, California, USA) (1988). 17. KIMME-SMITH, C, A review of mammography test objects for the calibration of resolution, contrast and exposure, Med. Phys., 16, 758-765 (1989). 18. OLSON, S L, BAHAA, W F, WINTER, P F ET AL, Breast calcifications: analysis of imaging properties, Radiology, 169, 329-332 (1988). 19. SWETS, J A and PICKET, R M, Evaluation of Diagnostic Systems: Methods from Signal Detection Theory (Academic Press, London) (1982). 20. WORKMAN, A and COWEN, A R, Exposure monitoring in photostimulable phosphor computed radiography, Radiat. Prot. Dosim., 43, 135-138 (1992). 21. NISHIKAWA, R M and YAFFE, M J, Signal-to-noise properties of mammographic film screen system, Med. Phys., 12, 32-39 (1985). 22. BUNCH, P C, HUFF, K E and VAN METTER, R, Analysis of the detective quantum efficiency of a radiographic screen-film combination, J. Opt. Soc. Am., 4, 902-909 (1987). 23. KUHN, H and KNUPFER, W, Imaging characteristics of different mammographic screens, Med. Phys., 19, 449^57 (1992). 24. LUBINSKY, A R, WHITING, B R and OWEN, J F, Storage phosphor system for computed radiography: optical effects and detective quantum efficiency (DQE), Proc. SPIE, 767, 161-111 (1987). 25. BUNCH, P C, Detective quantum efficiency of selected mammographic screen-film combinations, Proc. SPIE, 1090, 61-11 (1989). 26. WOLF, M, Signal-to-noise ratio and the detection of detail in non-white noise, Photogr. Sci. Eng., 24, 99-103 (1980). 995

A Workman, A R Cowen and D S Brettle 27. NAB, H W, KASSEMIER, N, VON ERRING, L J and HENDRIKS, J H, Comparison of digital and conventional mammography: an ROC study of 270 mammograms, Med. Inf., 17, 125-131 (1992). 28. KASSEMIER, N, FRIELENG, J T and HENDRIKS, J H, Spatial resolution in digital mammography, Invest. Radiol., 28, 413^19 (1993). 29. OESTMAN, J W, KOPANS, S D, HALL, D A ET AL, Equivalent diagnostic performance of digitised storage phosphor and conventional mammography in the detection of malignant microcalcifications, Invest. Radiol., 121, 9-17 (1989).

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