- Email: [email protected]
Digital mammography—detector considerations and new applications
Nuclear Instruments and Methods in Physics Research A 471 (2001) 6–11
Digital mammographyFdetector considerations and new applications Martin J. Yaffe* Imaging/Bioengineering Research, Sunnybrook & Women’s College Health Sciences Centre, Room S6-57, University of Toronto, 2075 Bayview Avenue, Toronto, Ontario, Canada M4N 3M5
Abstract Digital mammography offers the potential to improve the accuracy of detection and diagnosis of breast cancer. Key elements of the system are the X-ray detector and the display device and software. The design of each of these must be optimized to provide the expected performance of this new imaging technique. In this article, the motivation for digital mammography is briefly reviewed and the technical requirements of the X-ray detection system are discussed with reference to the current technology. Some important future applications of digital mammography are described. r 2001 Elsevier Science B.V. All rights reserved. Keywords: Digital mammography; X-ray detectors; Breast cancer; Mammography
1. Introduction Mammography is effective in imaging small breast tumours and indirect signs of cancer by recording changes in X-ray attenuation caused by variations in tissue composition of the breast. In conventional, film-based mammography, the screen-film system must act as the detector as well as the image storage and display device. Compromises are inevitable. Film is limited by its nonlinear (sigmoidal) response. In areas of the breast that are highly attenuating to X-rays and also in areas that are radiolucent, (i.e. where the transmitted X-ray fluence is low or high) the gradient of the film response is low, resulting in poor image contrast. In addition, the signal-to-noise properties
*Tel.: +1-416-480-5715; fax: +1-416-480-5714. E-mail address: martin.yaff[email protected] (M.J. Yaffe).
of the screen-film image receptor are far from optimal [1]. A major feature of digital mammography is that the processes of image acquisition, image display and image storage and retrieval are decoupled, allowing each to be optimized individually. The expected improvement in image quality offers potential advantages in terms of more accurate detection, diagnosis and image-guided treatment of disease. Another feature of digital imaging is the much greater ease of extracting useful quantitative information from the images. The performance of the detector is key in determining the efficiency with which images are produced and their intrinsic quality. This quality is likely to be closely correlated with the diagnostic value of the images. Some of the important features of detector performance are the quantum efficiency of the detector, its sensitivity to X-rays, its linearity, dynamic range and uniformity. High
0168-9002/01/$ - see front matter r 2001 Elsevier Science B.V. All rights reserved. PII: S 0 1 6 8 - 9 0 0 2 ( 0 1 ) 0 1 0 5 6 - 7
M.J. Yaffe / Nuclear Instruments and Methods in Physics Research A 471 (2001) 6–11
spatial resolution is particularly important in mammography, where small calcific particles and fine fibres radiating from a tumour must be seen. In addition, the geometrical configuration associated with the detector and its packaging must not impair access to imaging as much breast tissue as possible. In particular, there should not be dead or inaccessible areas of the detector near the patient’s chest wall.
7
will increase sharply at this energy. In addition, K X-ray fluorescence is likely to be produced in the detector after a photoelectric interaction. Escape of some fluorescent X-rays can cause variation in detector sensitivity, leading to increased noise [2]. Currently, the most common detector material for digital mammography is thallium-activated cesium iodide, CsI(Tl). The K edge of iodine is 33.2 keV, so that for most applications, X-ray interactions are by the L photoelectric effect and there is no K fluorescence produced.
2. Quantum interaction efficiency The probability of interaction or quantum efficiency for quanta of energy E ¼ hn is given by
3. Sensitivity
ZðEÞ ¼ 1 emðEÞT
Ultimately, the output from an X-ray detector used for digital radiography is an electrical signal. The sensitivity of the detector depends on the product of Z; the conversion efficiency of the detector and the efficiency of signal collection. We can define the conversion factor in terms of the charge or optical signal produced by the detector (before any external amplification) per interacting X-ray quantum of a specified energy. Typically, the conversion factor can be expressed as the energy, w; necessary to release a light photon in a phosphor, an electron–hole pair in a photoconductor (or semiconductor) or an electron–ion pair in a gaseous detector. Values of w for some detector materials used or considered for medical imaging are given in Table 1. Typically, this value is approximately three times higher the bandgap energy for the material.
ð1Þ
where m is the linear attenuation coefficient of the detector material and T is the active thickness of the detector. Diagnostic X-ray beams arise from a bremsstrahlung source and are, therefore, polyenergetic, so Z will vary across the spectrum used to produce an image. For mammography, accelerating voltages between 24 and 33 kV from Mo, Rh or W anode targets are typical, although higher voltages may be useful for some digital applications. The quantum interaction efficiency can be maximized by increasing the detector thickness or using detector materials of high atomic number or electron density. All currently-used detectors for digital mammography employ a phosphor which converts the X-ray energy to light. For these, an increase in detector thickness results in a reduction in spatial resolution due to lateral diffusion of light between the point of creation and collection. For such detectors, it may be necessary to compromise quantum interaction efficiency to maintain a required level of spatial resolution. For this reason, detectors of high Z and electron density, re ; are desirable. At diagnostic energies, the main mode of interaction, especially in high-Z detector materials, is by the photoelectric effect, so that there will be very little X-ray scattering within the detector. Generally, Z falls as energy increases. If the K absorption edge of the detector material falls within the energy spectrum of the X-ray beam, Z
4. Spatial resolution Detectors for digital radiography are often composed of discrete elements or dels, generally, of constant size and spacing. The dimension of the active portion of each detector element defines an aperture. The aperture determines the spatial frequency response of the detector. For example, if the aperture is square with dimension, d; then the modulation transfer function (MTF) of the detector will be of the form sinc ( f ), where f is the spatial frequency along the x- or y-directions, and the MTF will have its first zero at the frequency
8
M.J. Yaffe / Nuclear Instruments and Methods in Physics Research A 471 (2001) 6–11
Table 1 Some characteristics of phosphors and photoconductors used as X-ray detectors for digital radiography, including atomic number, Z; and K absorption energy, EK ; of the principal absorbing elements. Sensitivity is expressed as the energy, w; which must be absorbed to release a quantum of light in a phosphor or an electron-hole pair in a photoconductor Material
Z
EK (keV)
W (eV)
CdTe Crystalline Si PbI2 HgI2 Amorphous selenium CsI(Tl) Gd2O2S BaFBr (as photostim. phosphor)
48/52 14 82/53 80/53 34 55/53 64 56/35
26.7/31.8 1.8 88/33.2 83.1/33.2 12.7 36.0/33.2 50.2 37.4/13.5
4.4 3.6 5–8 [3] 13 [3] 50 (at 10 V/mm)a 19 13 50–100b
Some detectors are not pixellated at the X-ray absorption stage, but rather d and p are defined in their readout mechanism. This is the case for the photostimulable phosphor detector system described below, where the phosphor plate is continuous, but the laser readout samples the plate at discrete locations. This can provide some flexibility in independently setting the sampling interval (scanning raster) and effective aperture size (laser spot size) to avoid aliasing. Issues of sampling in digital radiographic systems have been reviewed by Dobbins [5]. The required sample interval for digital mammography is still a controversial issue, however, it is probably in the range 50–100 mm.
a
7 eV (theoretical value at infinite field). Estimated by multiplying the bandgap of 8.3 eV by 3 and then 2 for the 50% efficiency of trap filling during X-ray exposure. The higher value reflects a possible additional loss of up to a factor of two is due to retrapping during readout. b
f ¼ d 1 ; expressed in the plane of the detector. A detector with d¼ 50 mm will have an MTF with its first zero at f ¼ 20 cycles/mm. The sampling interval, p; of the detector, is the pitch between sensitive elements or measurements. The fill factor provides a measure of the detector’s geometric efficiency in capturing incident X-rays and is given by d 2 =p2 : The sampling theorem states that only spatial frequencies in the pattern below ð2pÞ1 (the Nyquist frequency) can be faithfully imaged. If the pattern contains higher frequencies, then a phenomenon known as aliasing occurs wherein the frequency spectrum of the image pattern beyond the Nyquist frequency is mirrored about that frequency and added to the spectrum of lower frequencies, increasing the apparent spectral content of the image at these lower frequencies [4]. In a detector composed of discrete elements, the smallest sampling interval in a single image acquisition is p ¼ d; so that the Nyquist frequency is ð2dÞ1 while the aperture response falls to zero at twice that frequency (higher if the dimension of the sensitive region of the detector element is smaller than d; e.g. because the fill factor of the detector element is less than 1.0).
5. Current detector systems Various types of X-ray detectors have been considered for digital mammography. The technologies which have been incorporated into clinical systems to date include: (a) CsI(Tl) coupled to CCD arrays through straight or demagnifying fibre-optic tapers, (b) CsI(Tl) deposited upon large-area arrays of photodiodes and readout switches fabricated on amorphous silicon and (c) photostimulable phosphors with laser readout. In addition, the detector can be a full-area or a scanning device. The full-area detector acquires the image with a single brief X-ray exposure. In scanning systems, the X-ray source is collimated to a narrow beam which is scanned across the breast in synchrony with a detector which measures the transmitted radiation. Scanning systems allow very efficient rejection of scattered radiation, improving contrast and signalto-noise ratio without a significant dose penalty. Because the detector only measures radiation transmitted through part of the breast at any one time, for a given del size, fewer detector elements are required than in a full-area system, thereby reducing detector cost. Scanning systems generally require longer total exposure time, necessitating higher X-ray tube heat loading. Because each part of the breast is exposed for only a very short time, there is no motion blurring, but if the breast moves during scanning, mis-registration artefacts could
M.J. Yaffe / Nuclear Instruments and Methods in Physics Research A 471 (2001) 6–11
result. Finally, the longer scan time restricts the image repetition rate and may be a limitation in dynamic studies. Most of the phosphor-based detectors for digital mammography employ CsI(Tl) as the X-ray absorber. The cesium iodide can be fabricated as pillar or needle-like structures which act like fibre optic channels, conducting light produced within the phosphor along the length of the needles, thereby inhibiting lateral spread of the light. This allows the detector to be made thick enough to at least partially mitigate the usual compromise between Z and spatial resolution. The first digital mammography imaging systems that were introduced were small format devices for guidance of stereotactic biopsy procedures [6]. The detector consists of a phosphor X-ray absorber on the surface of a square de-magnifying fibre optic taper. The taper couples the phosphor to an area CCD array. Demagnification by approximately a factor of two can be performed with acceptable efficiency. Greater demagnification leads to signal loss that reduces the SNR of the images [7]. Fig. 1(a) illustrates a full-area detector for digital mammography (Hologics Inc) that is fabricated as a mosaic of 3 4 of these detector modules. The dimension of an individual detector element (del) is approximately 40 mm. Another system, developed by Fischer Imaging Inc., employs a detector which is of a slot format. A CsI(Tl) phosphor is coupled through a fibre optic faceplate to a CCD (Fig. 1(b)) and the image is acquired by scanning the detector along an arc across the chest wall behind the breast. This arc lies on a circle centred at the focal spot of the X-ray tube. During image acquisition, the X-ray tube pivots about that point, with the X-rays collimated into a fan beam, matching the area of the detector. The X-ray tube is activated during the entire scan and as the detector and slot beam move across the breast, the charge produced in the elements of the CCD is shifted down CCD columns at an equal rate, but in the opposite direction, so that the charge signal resulting from X-rays transmitted through a particular path in the breast is integrated in the CCD. When each charge packet reaches the last row of the CCD, it is read out and
9
digitized. This technique is called time delay integration (TDI). In the Fischer system, the del dimension is approximately 50 mm. A third design employed by general electric incorporates a large-area matrix of photodiodes formed on an amorphous silicon plate (Fig. 1(c)). Each light-sensitive diode element is connected by a thin film transistor (TFT) switch to a series of control lines and data lines. Light produced in the CsI(Tl) absorber layer that lies over the readout plate produces charge on the photodiodes which is read out and digitized. In the configuration used for digital mammography, the del is 100 mm. The fourth system, manufactured by Fuji, employs a photo-stimulable X-ray absorbing phosphor material, typically BaFCl. The system (Fig. 1(d)) is similar in operation to that of the detectors that have been used for several years for computed radiography [8]. In response to absorption of X-rays, electronic charges are stored in ‘‘traps’’ in the crystalline material of the phosphor where they remain stable for some time. The number of traps filled is proportional to the amount of radiation incident on the phosphor. After exposure, the phosphor plate is removed and is read out by scanning with a fine helium–neon laser beam. The red light from the laser ‘‘discharges’’ the traps, causing stimulated emission of blue light. The blue light is collected and measured with a photo-multiplier tube and the resulting signal is digitized to form the image. Currently the effective del size, which is determined by the laser scanning parameters, is 100 mm.
6. Opportunities for new detectors While all of these detector technologies have been applied clinically and have shown improved contrast-detail performance compared to conventional screen-film mammography, there is certainly potential for further improvement of detectors. Considerable work is taking place on the development of direct-conversion materials for digital radiography. These are materials in which the absorption of X-rays
10
M.J. Yaffe / Nuclear Instruments and Methods in Physics Research A 471 (2001) 6–11
Fig. 1. Detector systems currently used for digital mammography: (a) mosaic system with phosphor coupled to CCD through fibre optic taper; (b) scanning system; (c) amorphous silicon flat-panel detector with CsI(Tl); (d) photostimulable phosphor detector.
directly yields charge which can be measured and digitized. Materials under investigation include amorphous selenium [9,10], cadmium zinc telluride [11], lead iodide [3], mercuric iodide [3] and others. By eliminating the intermediate stage of X-ray to light conversion, it is expected that the noise characteristics of detectors using these new materials can be improved beyond that possible with phosphors. Additionally, be-
cause these detectors produce their initial signal directly as electronic charge, which can be easily collected by means of an electric field, it is possible to minimize lateral spread of the signal, thereby opening the possibility of extremely high spatial resolution. More detail regarding detectors for digital radiography is included in a review article by Yaffe and Rowlands [12].
M.J. Yaffe / Nuclear Instruments and Methods in Physics Research A 471 (2001) 6–11
7. Applications of digital mammography As digital mammography matures, image quality is likely to improve further, contributing to greater diagnostic accuracy. Nevertheless, the technology is considerably more expensive than that used in conventional screen-film mammography. For this reason, the clinical acceptance of digital mammography will almost certainly be based not only on image quality, but also upon new applications based on digital mammography that will enhance its value to the physician. One important application, already well on its way, is computer-aided detection and diagnosis which uses computer algorithms (feature extraction, neural nets, etc.) to emulate a second reader of the mammograms to increase sensitivity and/or specificity [13,14]. Another is telemammography, which can make high quality mammographic interpretation accessible even to remote, sparsely populated areas where expert mammographic radiologists are not available [15]. Contrast mammography uses angiographic techniques with digital mammography to better outline the extent of disease for surgical planning and to demonstrate tumours that may not be seen on normal mammography. Stereoscopic and tomographic approaches are also much more easily accomplished when images are directly acquired in digital form. Finally, digital mammography may be useful in providing more direct intra-operative guidance of treatment.
References [1] R.M. Nishikawa, G.E. Mawdsley, A. Fenster, M.J. Yaffe, Med. Phys. 14 (1987) 717.
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[2] R. Swank, J. Appl. Phys. 44 (1973) 4199. [3] R.A. Street, S.E. Ready, J.T. Rahn, et al., High resolution, direct detection X-ray imagers, in: J.T. Dobbins III, J.M. Boone (Eds.), Medical Imaging 2000, Proceedings of SPIE3977, 2000, pp. 418–428. [4] J.S. Bendat, A.G. Piersol, Random Data Analysis and Measurement Techniques, 2nd Edition Wiley, New York, 1986, p. 338. [5] J.T. Dobbins III, Med. Phys. 22 (1995) 171. [6] H. Roehrig, T. Yu, W.V. Schempp, Proc. SPIE 2279 (1994) 388. [7] A.D.A. Maidment, M.J. Yaffe, Phys. Med. Biol. 41 (1996) 475. [8] J.T. Dobbins III, D.L. Ergun, L. Rutz, et al., Med. Phys. 22 (1995) 171. [9] J.A. Rowlands, J. Yorkston, Flat panel detectors for digital radiography, in: J. Beutel, H.L. Kundel, R.L. Van Metter (Eds.), Handbook of Medical Imaging, SPIE Press, 2000, pp. 223–328. [10] A. Debrie, B. Polischuk, H. Rougeot, et al., Quantitative analysis of performance of selenium flat-panel detector for interventional mammography, in: J.T. Dobbins III, J.M Boone (Eds.), Medical Imaging 2000, Vol. SPIE-3977, 2000, pp. 176–184. [11] J.G. Mainprize, N.L Ford, S.Yin , T.O. Tumer, E. Gordon, W.J. Hamilton, M.J.Yaffe, Semiconductor materials for digital mammography, in: J.T. Dobbins III, J.M Boone (Eds.), Medical Imaging 2000, Vol. SPIE-3977, 2000, pp. 152–158. [12] M.J. Yaffe, J.A. Rowlands, Phys. Med. Biol. 42 (1997) 1. [13] N. Karssemeijer, J.H. Hendriks, Eur. Radiol. 7 (1997) 743. [14] M.L. Giger, Z. Huo, M.A. Kupinski, C.J. Vyborny, Computer-aided diagnosis in mammography, in: M. Sonka, J.M. Fitzpatrick (Eds.), Handbook of Medical Imaging, Vol. 2, SPIE Press, 2000, pp. 915–1004. [15] S.L. Lou, E.A. Sickles, H.K. Huang, D. Hoogstrate, F. Cao, J. Wang, M. Jahangiri, IEEE Trans. Inf. Technol. Biomed. 1 (1997) 270.
Digital mammographyFdetector considerations and new applications Martin J. Yaffe* Imaging/Bioengineering Research, Sunnybrook & Women’s College Health Sciences Centre, Room S6-57, University of Toronto, 2075 Bayview Avenue, Toronto, Ontario, Canada M4N 3M5
Abstract Digital mammography offers the potential to improve the accuracy of detection and diagnosis of breast cancer. Key elements of the system are the X-ray detector and the display device and software. The design of each of these must be optimized to provide the expected performance of this new imaging technique. In this article, the motivation for digital mammography is briefly reviewed and the technical requirements of the X-ray detection system are discussed with reference to the current technology. Some important future applications of digital mammography are described. r 2001 Elsevier Science B.V. All rights reserved. Keywords: Digital mammography; X-ray detectors; Breast cancer; Mammography
1. Introduction Mammography is effective in imaging small breast tumours and indirect signs of cancer by recording changes in X-ray attenuation caused by variations in tissue composition of the breast. In conventional, film-based mammography, the screen-film system must act as the detector as well as the image storage and display device. Compromises are inevitable. Film is limited by its nonlinear (sigmoidal) response. In areas of the breast that are highly attenuating to X-rays and also in areas that are radiolucent, (i.e. where the transmitted X-ray fluence is low or high) the gradient of the film response is low, resulting in poor image contrast. In addition, the signal-to-noise properties
*Tel.: +1-416-480-5715; fax: +1-416-480-5714. E-mail address: martin.yaff[email protected] (M.J. Yaffe).
of the screen-film image receptor are far from optimal [1]. A major feature of digital mammography is that the processes of image acquisition, image display and image storage and retrieval are decoupled, allowing each to be optimized individually. The expected improvement in image quality offers potential advantages in terms of more accurate detection, diagnosis and image-guided treatment of disease. Another feature of digital imaging is the much greater ease of extracting useful quantitative information from the images. The performance of the detector is key in determining the efficiency with which images are produced and their intrinsic quality. This quality is likely to be closely correlated with the diagnostic value of the images. Some of the important features of detector performance are the quantum efficiency of the detector, its sensitivity to X-rays, its linearity, dynamic range and uniformity. High
0168-9002/01/$ - see front matter r 2001 Elsevier Science B.V. All rights reserved. PII: S 0 1 6 8 - 9 0 0 2 ( 0 1 ) 0 1 0 5 6 - 7
M.J. Yaffe / Nuclear Instruments and Methods in Physics Research A 471 (2001) 6–11
spatial resolution is particularly important in mammography, where small calcific particles and fine fibres radiating from a tumour must be seen. In addition, the geometrical configuration associated with the detector and its packaging must not impair access to imaging as much breast tissue as possible. In particular, there should not be dead or inaccessible areas of the detector near the patient’s chest wall.
7
will increase sharply at this energy. In addition, K X-ray fluorescence is likely to be produced in the detector after a photoelectric interaction. Escape of some fluorescent X-rays can cause variation in detector sensitivity, leading to increased noise [2]. Currently, the most common detector material for digital mammography is thallium-activated cesium iodide, CsI(Tl). The K edge of iodine is 33.2 keV, so that for most applications, X-ray interactions are by the L photoelectric effect and there is no K fluorescence produced.
2. Quantum interaction efficiency The probability of interaction or quantum efficiency for quanta of energy E ¼ hn is given by
3. Sensitivity
ZðEÞ ¼ 1 emðEÞT
Ultimately, the output from an X-ray detector used for digital radiography is an electrical signal. The sensitivity of the detector depends on the product of Z; the conversion efficiency of the detector and the efficiency of signal collection. We can define the conversion factor in terms of the charge or optical signal produced by the detector (before any external amplification) per interacting X-ray quantum of a specified energy. Typically, the conversion factor can be expressed as the energy, w; necessary to release a light photon in a phosphor, an electron–hole pair in a photoconductor (or semiconductor) or an electron–ion pair in a gaseous detector. Values of w for some detector materials used or considered for medical imaging are given in Table 1. Typically, this value is approximately three times higher the bandgap energy for the material.
ð1Þ
where m is the linear attenuation coefficient of the detector material and T is the active thickness of the detector. Diagnostic X-ray beams arise from a bremsstrahlung source and are, therefore, polyenergetic, so Z will vary across the spectrum used to produce an image. For mammography, accelerating voltages between 24 and 33 kV from Mo, Rh or W anode targets are typical, although higher voltages may be useful for some digital applications. The quantum interaction efficiency can be maximized by increasing the detector thickness or using detector materials of high atomic number or electron density. All currently-used detectors for digital mammography employ a phosphor which converts the X-ray energy to light. For these, an increase in detector thickness results in a reduction in spatial resolution due to lateral diffusion of light between the point of creation and collection. For such detectors, it may be necessary to compromise quantum interaction efficiency to maintain a required level of spatial resolution. For this reason, detectors of high Z and electron density, re ; are desirable. At diagnostic energies, the main mode of interaction, especially in high-Z detector materials, is by the photoelectric effect, so that there will be very little X-ray scattering within the detector. Generally, Z falls as energy increases. If the K absorption edge of the detector material falls within the energy spectrum of the X-ray beam, Z
4. Spatial resolution Detectors for digital radiography are often composed of discrete elements or dels, generally, of constant size and spacing. The dimension of the active portion of each detector element defines an aperture. The aperture determines the spatial frequency response of the detector. For example, if the aperture is square with dimension, d; then the modulation transfer function (MTF) of the detector will be of the form sinc ( f ), where f is the spatial frequency along the x- or y-directions, and the MTF will have its first zero at the frequency
8
M.J. Yaffe / Nuclear Instruments and Methods in Physics Research A 471 (2001) 6–11
Table 1 Some characteristics of phosphors and photoconductors used as X-ray detectors for digital radiography, including atomic number, Z; and K absorption energy, EK ; of the principal absorbing elements. Sensitivity is expressed as the energy, w; which must be absorbed to release a quantum of light in a phosphor or an electron-hole pair in a photoconductor Material
Z
EK (keV)
W (eV)
CdTe Crystalline Si PbI2 HgI2 Amorphous selenium CsI(Tl) Gd2O2S BaFBr (as photostim. phosphor)
48/52 14 82/53 80/53 34 55/53 64 56/35
26.7/31.8 1.8 88/33.2 83.1/33.2 12.7 36.0/33.2 50.2 37.4/13.5
4.4 3.6 5–8 [3] 13 [3] 50 (at 10 V/mm)a 19 13 50–100b
Some detectors are not pixellated at the X-ray absorption stage, but rather d and p are defined in their readout mechanism. This is the case for the photostimulable phosphor detector system described below, where the phosphor plate is continuous, but the laser readout samples the plate at discrete locations. This can provide some flexibility in independently setting the sampling interval (scanning raster) and effective aperture size (laser spot size) to avoid aliasing. Issues of sampling in digital radiographic systems have been reviewed by Dobbins [5]. The required sample interval for digital mammography is still a controversial issue, however, it is probably in the range 50–100 mm.
a
7 eV (theoretical value at infinite field). Estimated by multiplying the bandgap of 8.3 eV by 3 and then 2 for the 50% efficiency of trap filling during X-ray exposure. The higher value reflects a possible additional loss of up to a factor of two is due to retrapping during readout. b
f ¼ d 1 ; expressed in the plane of the detector. A detector with d¼ 50 mm will have an MTF with its first zero at f ¼ 20 cycles/mm. The sampling interval, p; of the detector, is the pitch between sensitive elements or measurements. The fill factor provides a measure of the detector’s geometric efficiency in capturing incident X-rays and is given by d 2 =p2 : The sampling theorem states that only spatial frequencies in the pattern below ð2pÞ1 (the Nyquist frequency) can be faithfully imaged. If the pattern contains higher frequencies, then a phenomenon known as aliasing occurs wherein the frequency spectrum of the image pattern beyond the Nyquist frequency is mirrored about that frequency and added to the spectrum of lower frequencies, increasing the apparent spectral content of the image at these lower frequencies [4]. In a detector composed of discrete elements, the smallest sampling interval in a single image acquisition is p ¼ d; so that the Nyquist frequency is ð2dÞ1 while the aperture response falls to zero at twice that frequency (higher if the dimension of the sensitive region of the detector element is smaller than d; e.g. because the fill factor of the detector element is less than 1.0).
5. Current detector systems Various types of X-ray detectors have been considered for digital mammography. The technologies which have been incorporated into clinical systems to date include: (a) CsI(Tl) coupled to CCD arrays through straight or demagnifying fibre-optic tapers, (b) CsI(Tl) deposited upon large-area arrays of photodiodes and readout switches fabricated on amorphous silicon and (c) photostimulable phosphors with laser readout. In addition, the detector can be a full-area or a scanning device. The full-area detector acquires the image with a single brief X-ray exposure. In scanning systems, the X-ray source is collimated to a narrow beam which is scanned across the breast in synchrony with a detector which measures the transmitted radiation. Scanning systems allow very efficient rejection of scattered radiation, improving contrast and signalto-noise ratio without a significant dose penalty. Because the detector only measures radiation transmitted through part of the breast at any one time, for a given del size, fewer detector elements are required than in a full-area system, thereby reducing detector cost. Scanning systems generally require longer total exposure time, necessitating higher X-ray tube heat loading. Because each part of the breast is exposed for only a very short time, there is no motion blurring, but if the breast moves during scanning, mis-registration artefacts could
M.J. Yaffe / Nuclear Instruments and Methods in Physics Research A 471 (2001) 6–11
result. Finally, the longer scan time restricts the image repetition rate and may be a limitation in dynamic studies. Most of the phosphor-based detectors for digital mammography employ CsI(Tl) as the X-ray absorber. The cesium iodide can be fabricated as pillar or needle-like structures which act like fibre optic channels, conducting light produced within the phosphor along the length of the needles, thereby inhibiting lateral spread of the light. This allows the detector to be made thick enough to at least partially mitigate the usual compromise between Z and spatial resolution. The first digital mammography imaging systems that were introduced were small format devices for guidance of stereotactic biopsy procedures [6]. The detector consists of a phosphor X-ray absorber on the surface of a square de-magnifying fibre optic taper. The taper couples the phosphor to an area CCD array. Demagnification by approximately a factor of two can be performed with acceptable efficiency. Greater demagnification leads to signal loss that reduces the SNR of the images [7]. Fig. 1(a) illustrates a full-area detector for digital mammography (Hologics Inc) that is fabricated as a mosaic of 3 4 of these detector modules. The dimension of an individual detector element (del) is approximately 40 mm. Another system, developed by Fischer Imaging Inc., employs a detector which is of a slot format. A CsI(Tl) phosphor is coupled through a fibre optic faceplate to a CCD (Fig. 1(b)) and the image is acquired by scanning the detector along an arc across the chest wall behind the breast. This arc lies on a circle centred at the focal spot of the X-ray tube. During image acquisition, the X-ray tube pivots about that point, with the X-rays collimated into a fan beam, matching the area of the detector. The X-ray tube is activated during the entire scan and as the detector and slot beam move across the breast, the charge produced in the elements of the CCD is shifted down CCD columns at an equal rate, but in the opposite direction, so that the charge signal resulting from X-rays transmitted through a particular path in the breast is integrated in the CCD. When each charge packet reaches the last row of the CCD, it is read out and
9
digitized. This technique is called time delay integration (TDI). In the Fischer system, the del dimension is approximately 50 mm. A third design employed by general electric incorporates a large-area matrix of photodiodes formed on an amorphous silicon plate (Fig. 1(c)). Each light-sensitive diode element is connected by a thin film transistor (TFT) switch to a series of control lines and data lines. Light produced in the CsI(Tl) absorber layer that lies over the readout plate produces charge on the photodiodes which is read out and digitized. In the configuration used for digital mammography, the del is 100 mm. The fourth system, manufactured by Fuji, employs a photo-stimulable X-ray absorbing phosphor material, typically BaFCl. The system (Fig. 1(d)) is similar in operation to that of the detectors that have been used for several years for computed radiography [8]. In response to absorption of X-rays, electronic charges are stored in ‘‘traps’’ in the crystalline material of the phosphor where they remain stable for some time. The number of traps filled is proportional to the amount of radiation incident on the phosphor. After exposure, the phosphor plate is removed and is read out by scanning with a fine helium–neon laser beam. The red light from the laser ‘‘discharges’’ the traps, causing stimulated emission of blue light. The blue light is collected and measured with a photo-multiplier tube and the resulting signal is digitized to form the image. Currently the effective del size, which is determined by the laser scanning parameters, is 100 mm.
6. Opportunities for new detectors While all of these detector technologies have been applied clinically and have shown improved contrast-detail performance compared to conventional screen-film mammography, there is certainly potential for further improvement of detectors. Considerable work is taking place on the development of direct-conversion materials for digital radiography. These are materials in which the absorption of X-rays
10
M.J. Yaffe / Nuclear Instruments and Methods in Physics Research A 471 (2001) 6–11
Fig. 1. Detector systems currently used for digital mammography: (a) mosaic system with phosphor coupled to CCD through fibre optic taper; (b) scanning system; (c) amorphous silicon flat-panel detector with CsI(Tl); (d) photostimulable phosphor detector.
directly yields charge which can be measured and digitized. Materials under investigation include amorphous selenium [9,10], cadmium zinc telluride [11], lead iodide [3], mercuric iodide [3] and others. By eliminating the intermediate stage of X-ray to light conversion, it is expected that the noise characteristics of detectors using these new materials can be improved beyond that possible with phosphors. Additionally, be-
cause these detectors produce their initial signal directly as electronic charge, which can be easily collected by means of an electric field, it is possible to minimize lateral spread of the signal, thereby opening the possibility of extremely high spatial resolution. More detail regarding detectors for digital radiography is included in a review article by Yaffe and Rowlands [12].
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7. Applications of digital mammography As digital mammography matures, image quality is likely to improve further, contributing to greater diagnostic accuracy. Nevertheless, the technology is considerably more expensive than that used in conventional screen-film mammography. For this reason, the clinical acceptance of digital mammography will almost certainly be based not only on image quality, but also upon new applications based on digital mammography that will enhance its value to the physician. One important application, already well on its way, is computer-aided detection and diagnosis which uses computer algorithms (feature extraction, neural nets, etc.) to emulate a second reader of the mammograms to increase sensitivity and/or specificity [13,14]. Another is telemammography, which can make high quality mammographic interpretation accessible even to remote, sparsely populated areas where expert mammographic radiologists are not available [15]. Contrast mammography uses angiographic techniques with digital mammography to better outline the extent of disease for surgical planning and to demonstrate tumours that may not be seen on normal mammography. Stereoscopic and tomographic approaches are also much more easily accomplished when images are directly acquired in digital form. Finally, digital mammography may be useful in providing more direct intra-operative guidance of treatment.
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