- Email: [email protected]
Experience with scintillators for PET: towards the fifth generation of PET scanners
ARTICLE IN PRESS
Nuclear Instruments and Methods in Physics Research A 525 (2004) 242–248
Experience with scintillators for PET: towards the fifth generation of PET scanners L. Erikssona,b,*, D. Townsendc, M. Erikssona,b, C. Melcherd, M. Schmanda, B. Bendriema, R. Nuttd a
CPS Innovations, 810 Innovation Dr, Knoxville, TN 37932, USA b Karolinska Institute, Stockholm, Sweden c Department of Medicine and Radiology, University of Tennessee Medical Center, TN, USA d CTI Molecular Imaging, Knoxville TN, USA
Abstract Since the ECAT 2, the first commercial positron emission tomograph developed by EG&G ORTEC, four generations of scanners can be identified. The first such scanners were based on sodium iodide (NaI(Tl)) scintillators, although as early as 1978 the transition to bismuth germanate (BGO) detectors had begun. By 1981, second-generation PET scanners with up to four rings of BGO detectors were available commercially. The BGO block detector appeared in 1985, initiating the third generation of PET scanners with the potential to increase the axial coverage in a cost-effective manner. As with the second generation, the third generation of PET scanners incorporated lead septa to collimate the annihilation photons within transverse planes, thereby reducing the acquisition of scattered and random coincidences and limiting detector dead time. The fourth generation of PET scanners offered up to 15 cm axial coverage and incorporated retractable septa that permitted both 2D and 3D acquisition within the same scanner. Therefore, fifthgeneration scanner should be a fully 3D system with no septa, a 25–30 cm axial field-of-view, and a spatial resolution approaching the limits set by the physics of positron emission. Rather than using septa, limitation of randoms and scatter should be achieved directly at the detector by using a scintillator with high-light output, good energy resolution, and a fast scintillation decay time. The fast decay time will ensure low deadtime. The recent development of lutetium oxyorthosilicate (LSO), a scintillator with the required properties, suggests that a fifth generation of positron emission tomographs can now be attained. r 2004 Elsevier B.V. All rights reserved. PACS: 87.59.Vb; 87.62.+n Keywords: PET; Lutetium oxyorthosilicate; LSO; Whole-body imaging; 3D imaging; 3D data acquisition
1. Introduction
*Corresponding author. CPS Innovations, 810 Innovation Dr, Knoxville, TN 37932, USA Tel.: +1-865-218-2246; fax: +1-865-218-3000. E-mail address: [email protected] (L. Eriksson).
Positron emission tomography (PET) scanners have been available since the designs of the mid1970s that originated with the development of the PETT III [1] at Washington University in St. Louis. This initial design led to the ECAT 2
0168-9002/$ - see front matter r 2004 Elsevier B.V. All rights reserved. doi:10.1016/j.nima.2004.03.067
ARTICLE IN PRESS L. Eriksson et al. / Nuclear Instruments and Methods in Physics Research A 525 (2004) 242–248
scanner manufactured by EG&G ORTEC in Oak Ridge, Tennessee. These early designs were the forerunners of the high-performance PET technology available today. Over the past 25 years, PET scanner technology has undergone continuous and steady improvement, especially in the past 5 years since the introduction of reimbursement for PET imaging in a number of different cancers. By reviewing the progress of PET technology, four different generations can be identified even though the boundaries between consecutive generations may not always be well defined. This paper will describe the characteristics of each of the four generations and anticipate the design features of a fifth generation of PET scanner technology.
2. Four generations of pet scanners The first generation of PET scanners incorporated a single ring, or hexagonal array, of sodium iodide (NaI(Tl)) detectors. This scintillator had first been used in gamma cameras and was a natural choice for PET. However, while being an optimal scintillator at 140 keV for use in standard nuclear medicine with 99mTc, sodium iodide was less well suited to detect the higher energy, 511 keV photons from positron annihilation. A new scintillator, bismuth germanate (BGO) discovered in the early 1970s [2], had much higher sensitivity than sodium iodide at 511 keV, even though the light output was around 15% that of NaI(Tl). The potential of BGO for PET was first recognized and described in a paper by Cho et al. in 1977 [3]. The introduction of BGO gave rise to the second generation of PET scanners, initially incorporating only one or two rings of detectors. Designs with two or more detector rings, such as the Scanditronix PC-384-7b scanner, also included annular lead shields, or septa, between the detector rings. The septa served to shield the detectors from outof-plane scatter and randoms and reduce dead time by restricting coincidences to within the transaxial plane. In 1985, Casey and Nutt at CTI (Knoxville, TN) proposed the BGO detector block [4] opening up the possibility for multiple rings of detectors with small individual crystals covering an increased axial extent at a reasonable cost. The
243
block detector scanners of the mid-1980s such as the CTI 831 and the Scandotronix PC2048-15B represented the third generation of PET scanners. These designs still included septa between the detector rings to limit scatter and random coincidences and the axial coverage was either 5 cm or 10 cm depending on the number of rings of detector blocks (one or two) actually incorporated into the scanner. The fourth generation of scanners saw the introduction of retractable septa that, when fully retracted from the imaging volume, allowed coincidences to be collected between any pair of detector rings. This a 3D mode of operation increased the system sensitivity by factors of 5 or more, although scatter and randoms rates also increased. Such 3D data sets demanded a fully 3D reconstruction algorithm, the feasibility of which had earlier been demonstrated by Townsend et al. [5] by physically removing the septa from a thirdgeneration PET scanner. The first scanner with retractable septa, the ECAT 953B, primarily a brain scanner with a 35 cm patient port and 10 cm axial coverage, appeared in 1990. Results and comparisons between 2D and 3D operation of the ECAT 953B were first published in Ref [6]. By 1991, all septa-retractable fourth generation scanners incorporated an additional ring of detector blocks, thus covering 15 cm axially. Examples of fourth-generation PET scanners include the ECAT EXACT and the ECAT EXACT HR+ from CTI PET Systems (Knoxville, TN) and the ADVANCE scanner from GE Medical Systems (Milwaukee, WI). CTI PET Systems also designed the unique ECAT EXACT HR++, a PET scanner that incorporated no septa and covered an axial field of 24 cm with 48 rings of 4 mm BGO crystals. The example of this design was installed at Hammersmith Hospital, London in 1996 [7].
3. Towards a fifth generation of PET scanners A fifth generation of scanners will obviously improve and extend the performance of the fourth generation, just as the fourth generation represented an improvement over the third. Parameters to be improved include sensitivity, spatial
ARTICLE IN PRESS 244
L. Eriksson et al. / Nuclear Instruments and Methods in Physics Research A 525 (2004) 242–248
resolution, axial coverage, random rates and scatter fraction, count rate and dead time. In general terms, the next generation should have increased axial coverage of at least 30 cm to make maximum use of the emitted photon flux, operate in 3D only (no septa) for maximum sensitivity, but comprising small detector elements of 4 mm or less. High spatial resolution is essential, particularly for imaging the brain, and, as with animal PET scanners, the resolution should be increased to the limit set by the physics of positron emission. To limit the levels of random and scattered coincidences, the detector must have good coincidence timing and energy resolution, respectively. A time resolution of 1–2 ns, or better, is desirable, with an energy resolution of at least 14%. For optimal count rate performance, a low dead time is required such that the processing time for detected photons is within 100–200 ns including time and energy validation and reset time. Depending on the size of the detector units and the incoming photon flux, the processing and reset time should be reduced below 200 ns. The limitations of spatial resolution with the current block detector design are well understood. The block uses Anger logic to localize the incident photon to a particular detector element. Spatial resolution is therefore limited by the size of the detector element (pixel) in relation to the size and sampling distance of the optical sensors such as the phototubes. The maximum number of pixels per optical sensor depends strongly on the sharing of the light, the design of the light guide, and the amount of light emitted from the scintillator by conversion of a 511 keV photon. With current PET scanner technology for imaging patients, typical pixel sizes are 6 6 d mm3 down to 2 2 d mm3, where the depth d is in the range 20–30 mm. It is well known that other factors also limit spatial resolution, such as the range of the positrons after emission from the nucleus and before annihilation, and the acolinearity of the photon pairs due to the non-zero momentum of the annihilating electron–positron system. The uncertainty due to positron range is a function of the energy of the positron when ejected from the nucleus, which in turn depends upon the particular positron-emitting isotope used. The uncertainty
due to acolinearity depends on the distance between the coincident detectors. For a whole body PET scanner with a ring diameter of 80 cm, the contribution to the spatial resolution from acolinearity and positron range (for an isotope such as 18F) will total about 1.8 mm (full-width at half-maximum). To match this uncertainty due to positron physics, detector pixels should be around 2–3 mm in size. The energy resolution of the scintillator influences the fraction of scattered coincidences collected during an acquisition. Since all scintillators have a finite energy resolution, the energy of the 511 keV incident photon estimated from the total output of the optical sensors will have a range of values around the photo peak. For a 15% energy resolution, this range will be about 400–650 keV; the scanner is operated with a lower level discriminator (LLD) setting of 400 keV without loss of photo-peak sensitivity. With better energy resolution, say 12%, the LLD can be raised to 450 keV without loss of photo-peak events. Measured with a 70 cm long, 20 cm diameter polyethylene cylinder with a line source insert (NEMA NU 2-2001), the scatter fraction for an energy window of 400–650 keV is around 40%, decreasing to 30% or less with an energy window of 450– 650 keV. The goal, therefore, is to identify a scintillator with the appropriate properties to match the requirements of the fifth generation, high-performance PET scanner. The scintillator should have high light output to meet the positioning accuracy required, good energy resolution, high density and photo fraction for good sensitivity, and fast rise time and decay to ensure good coincidence timing and low detector dead time. Despite over two decades as the PET scintillator of choice, BGO is no longer a good match for these demands. The low light output (15% of NaI) and long decay time of the scintillation (300 ns) limits the timing resolution to 5–6 ns and the light signal would require integration for 500–600 ns to ensure adequate energy resolution; a long integration time results in high detector dead time. There are a number of scintillators that are possible candidates for the fifth generation of highperformance PET scanners, as summarized in
ARTICLE IN PRESS L. Eriksson et al. / Nuclear Instruments and Methods in Physics Research A 525 (2004) 242–248
Table 1 compared with BGO and NaI(Tl). One candidate, gadolinium oxyorthosilicate (Gd2SiO5:Ce, or GSO) was first discovered in 1983 [8] and used in early PET scanners in conjunction with BGO. Two layer crystals, one layer of BGO and one of GSO were positioned on the same photomultiplier and pulse shape discrimination used to separate the two scintillators [9]. Unfortunately, GSO cleaves rather easily when cut or machined so special techniques have been developed to avoid the crystals fracturing during detector construction. GSO has recently been used in a brain PET scanner, the G-PET [10], and in a commercial whole-body PET scanner, the ALLEGRO from Phillips Medical Systems. The G-PET and ALLEGRO can be considered examples of the fifth generation of PET scanners. The Crystal Clear collaboration has explored the use of cerium-activated lutetium–yttrium– ortho-aluminate (LuxY1xAlO3:Ce or LuYAP) for PET, where typically x ¼ 0:720:8 [11]. The crystal has a number of promising features such as good stopping power and a fast decay component of around 26 ns. However, for LuYAP light output is comparable to BGO, while for lutetium orthoaluminate (LuAP) it is around twice that of BGO. In either case the low light output will limit the crystal identification accuracy of the Anger logic as it does with BGO. The disadvantage of low light output is overcome by the use of the scintillator lutetium oxyorthosilicate (Lu2SiO5:Ce, or LSO), the best detector for PET known at the present time [12]. LSO combines high density and photo fraction to give good stopping power for 511 keV photons, with a light output of 75% of NaI(Tl) that is more than double that of LuAP; the
245
scintillation decay time is around 40 ns. The crystal structure is rugged and easy to machine into pixels of different dimensions, and is non-hygroscopic that facilitates packaging. LSO is an obvious choice for the fifth generation of PET scanners. The first whole-body clinical PET scanner based on LSO detectors is the ECAT ACCEL (CPS Innovations, Knoxville, TN), a design similar to the BGO-based ECAT EXACT. The relative performance of the scintillators can be assessed by comparing the count rate characteristics of scanners comprising different crystals but with similar geometric design. One metric of count rate performance is the Noise Equivalent Count Rate (NECR) [13] defined by NECR ¼
½ðT þ SÞð1 scfÞ2 T þSþR
ð1Þ
In this equation T, S and R are the true, scattered and random coincidence rates, respectively. To compare different scintillators, a fixed geometry has been used based on the ECAT EXACT design: three rings of blocks, 48 blocks per ring, a block size of 52 mm 52 mm with a depth of 25 mm; in the ECAT EXACT the BGO crystal depth is actually 20 mm. The comparison is made for the NEMA NU-2 phantom (2002), a polyethylene cylinder 20 cm in diameter and 70 cm in length with an off-center line source insert. The results of the comparison for this phantom for different scintillators are shown in Fig. 1, clearly demonstrating the superiority of LSO as the detector material for PET. The NEC curves are generated with measured sensitivity values for the different scintillators and a time window of 3 ns for LSO
Table 1 Physical properties of different scintillators for PET
Density (g/cc) Effective Z Radiation length for 511 keV (cm) Decay time (ns) Photons/Mev Light yield (% NaI) Energy resolution 662 keV in % Hygroscopic
NaI
BGO
GSO
LSO
LuYAP
3.67 50.8 2.6 230 38,000 100 7 yes
7.13 75.2 1.12 300 8200 15 10 no
6.71 59.4 1.4 30–60 10,000 25 9 no
7.4 66.4 1.14 40–43 28,000 75 9 no
7.1 61.3 1.4 20–26, 200 6000–8000 B15 8 no
ARTICLE IN PRESS 246
L. Eriksson et al. / Nuclear Instruments and Methods in Physics Research A 525 (2004) 242–248
Fig 1. NEC comparison for different scintillators with a 450 keV LLD setting based on the ECAT EXACT geometry. Only detector dead time has been modelled.
and LuYAP, 4 ns for GSO and NaI(Tl) and 11 ns for BGO. BGO has the highest sensitivity at 511 keV of all the scintillators listed in Table 1. The dead time characteristics of a BGO scanner could be substantially improved by using smaller block sizes, although at the expense of an increased number of photo detectors and more complex electronics resulting in higher production costs. The main disadvantages to continuing to use BGO for the fifth-generation PET scanners are the low light output and long scintillation time. The time resolution of a pair of detectors is given approximately by the ratio of the scintillation time to the square root of the number of photo electrons created in the photomultiplier. The best timing resolution for BGO scanners is around 5–6 ns, implying a time window of 10–12 ns to recover full sensitivity. For a fifth-generation scanner, the time window should be limited only by the size of the transverse field-of-view; a time window of 3–4 ns would cover all time-of-flight variations. The timing goal should therefore be to achieve a subnanosecond level. The next step is to identify the optimal detector configuration for the fifth-generation PET scanner. An increase in axial length could be achieved by adding extra rings of blocks to the current three or four block systems. However, while the block design has been extremely successful for almost
two decades, it may be time to consider alternatives such as multiple large flat panels that would surround the patient in a hexagonal or octagonal configuration. In such LSO-based panels, a single large (5 cm) phototube can be shared by many pixels, thus resulting in significant cost saving by reducing the overall number of phototubes. While such a design may be cost-effective, a continuous light guide (between the phototubes and the scintillator) over such a large area may lead to pixel identification errors due to pile-up effects. However, it has been found that elaborate pile-up correction techniques may be used with only a minor increase in system dead time. As a geometrical configuration, it has been shown [14] that, for a given amount of detector material, there are some advantages to arrays with large axial extent covering a limited angular range transversely as opposed to complete transverse coverage over a more limited axial extent. Obviously the advantages and disadvantages of smaller blocks as opposed to larger panels need to be carefully reviewed. An estimate for the spatial resolution (s) of such a design can be obtained from the following expression (in mm) [15]: qffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffi s ¼ 1:25 ½ðd=22 þ ð0:0022DÞ2 þ r2 þ b2 þ ðdsÞ2 ð2Þ where d is the size of the detector element (pixel), D is the separation of the detectors, r is the range of the positron and b is a measure of the crystal identification uncertainty, taken as 1 mm for LSO. A linear sampling term (ds), expressed in units of d; has also been included in Eq. (2), where s p 0.5. The results yielded by this expression for a scanner of diameter 85 cm and pixel sizes of 2, 3, and 4 mm are shown in Table 2; the positron range contribution has been estimated for imaging with isotopes 18 F and 15O. The overall sensitivity of the scanner depends not only on the physical performance of the scintillator material but also on the effective coverage of the volume around the patient into which the radiation is emitted. This coverage is defined by the detector geometry and the packing fraction, the density of pixels per detector module.
ARTICLE IN PRESS L. Eriksson et al. / Nuclear Instruments and Methods in Physics Research A 525 (2004) 242–248
The packing fraction is defined as the square of the ratio of the active crystal volume to the total detector volume. For small, 2 or 3 mm pixels, the reflective surfaces that guide the light from the scintillator into the photo sensor must be thin. The results of a calculation of the packing fraction for different pixel sizes and different reflective surfaces are shown in Fig. 2. For 2 mm pixels, a scanner resolution around 3 mm is feasible for 18F, degrading to 4 mm with 15O due to the higher energy positron. For 3 mm pixels, the estimates become 3.5 mm for 18F and 4.5 mm for 15O, values which probably represent the best spatial resoluTable 2 Estimates of reconstructed image resolution in a 85 cm fifthgeneration system as a function of detector widths and isotope used Isotope
18
15
F
O
Linear sampling factor
Detector widths (mm) 2
3
4
s¼0 s ¼ 0:25 s ¼ 0:5
3.04 3.11 3.29
3.40 3.52 3.88
3.82 4.02 4.56
s¼0 s ¼ 0:25 s ¼ 0:5
4.04 4.09 4.23
4.44 4.54 4.82
4.87 5.02 5.47
The positron ranges for 18F and 15O has been calculated using the expressions by Levin et al [16].
Fig 2. Packing fractions in coincidence for a block size of 52 52 20 mm3 block as a function of different sizes of detector elements.
247
tion that can be achieved with a whole body scanner for oncology applications. Ever since the development of the gamma camera, the principal photo detector in medical imaging has been the photomultiplier tube (PMT). For PET, the use of solid state detectors such as the avalanche photodiode (APD) has been hampered, by among other things, the low level of light emission from BGO. This situation has changed with the significantly higher light emission from LSO. PMTs, however, are faster than APDs and with the built-in amplification in the phototube itself, an inexpensive preamplifier can be used. With LSO, the ratio between the spacing and the surface area of the PMTs relative to the pixel size is in the order of 10:1 and thus 2 mm pixels can be easily identified with a 19 mm PMT. This empirical limit could be taken slightly further by the use of a slotted light guide that constrains the light into predefined regions. Since the cost of a PMT is independent of the size, large panels of 2 mm 2 mm pixels will be expensive. APDs are less bulky, but slower and require high-performance pre-amplifiers. APDs would also work in a magnetic field thus making a combined PET/MR scanner design feasible for human applications. Their quantum efficiency is around three times higher than that of phototubes and single 5 mm 5 mm APDs are available from Hamamatsu (S8664-55) with good performance characteristics. The overall package, however, is 10 mm 10 mm in size, so a high packing fraction for a large detector array would be impossible at this time. There may also be stability issues with an
Fig 3. (a) Position profile taken for one block of an HRRT panel, (b) position profile taken with four Hamamatsu APDs.
ARTICLE IN PRESS 248
L. Eriksson et al. / Nuclear Instruments and Methods in Physics Research A 525 (2004) 242–248
APD system requiring strict temperature control of the array to minimize drifts. The positioning profile for a PMT-based block is shown in Fig. 3a compared with an APD-based readout using four Hamamatsu APDs in Fig. 3b. While APDs offer considerable potential, more work is needed before the APD will routinely replace the phototube in PET scanners.
physics of positron annihilation. It remains for these, or comparable designs to bring high resolution, high-sensitivity whole-body imaging to in oncology. This is the challenge of the future in PET instrumentation.
References 4. A fifth-generation PET scanner for oncology From the above discussion it is evident that a fifth-generation whole-body PET scanner based on pixel technology and with a reconstructed spatial resolution around 3 mm is feasible. At the present time, this can only be achieved with LSO, and for cost-effectiveness, large area panels may be the solution. APDs could be an alternative to PMTs within the next 5–6 years. Finally, an example of an operational, fifth generation PET scanner is the High-Resolution Research Tomograph (HRRT), developed by CPS Innovations (Knoxville). This design comprises eight panels of 9 13 blocks viewed by a 10 14 array of 19 mm diameter PMTs; each block is 19 mm 19 mm in size and consists of two layers of different scintillators, a 10 mm layer of LSO and a 10 mm layer of LYSO, cut into 8 8 pixels per block. The layers are identified by pulse shape discrimination thus reducing the degradation of spatial resolution due to depth-of-interaction. The effective diameter of the HRRT, a scanner designed primarily for brain imaging, is around 46 cm; for a 20 cm diameter transverse field-of-view, a measured spatial resolution of 2.5 mm can be achieved with 18 F-labeled tracers. While fifth-generation PET scanner designs such as the ALLEGRO and the ACCEL are already operational in the clinical arena, it is evident that they do not attain the limits of possibility with the new scintillator materials. For specialized applications such as the brain, designs such as the G-PET and the HRRT are reaching the limits set by the
[1] M.E. Phelps, E. Hoffman, N. Mullani, C. Higgins, M. Ter-Pogossian, IEEE Trans. Biomed. Eng. NS-23 (1976) 516. [2] M.J. Weber, R.R. Monchamp, J. Appl. Phys. 44 (1973) 5495. [3] Z.H. Cho, M.R. Farukhi, J. Nucl. Med. 18 (1977) 840. [4] M. Casey, R. Nutt, IEEE Trans. Nucl. Sci. NS-33 (1986) 760. [5] D. Townsend, T. Spinks, T. Jones, A. Geissbuhler, M. Defrise, M-C. Gilardi, J. Heather, Eur. J. Nucl. Med. 15 (1989) 741. [6] D. Townsend, A. Geissbuhler, M. Defrise, E.J. Hoffman, T.J. Spinks, D.L. Bailey, M.C. Gilardi, T. Jones, IEEE Trans. Med. Imaging 10 (4) (1991) 505. [7] T. Jones, D.L. Bailey, P.M. Bloomfield, T.J. Spinks, W.F. Jones, K. Vaigneur, J. Reed, J. Young, D.F. Newport, C. Moyers, M.E. Casey, R. Nutt, J. Nucl. Med. 37 (5) (1996) 85P. [8] K. Takagi, T. Fukazawa, Appl. Phys. Lett. 42 (1983) 43. [9] S. Holte, H. Ostertag, M. Kesselberg, J. Comput. Assisted Tomogr. 11 (1987) 691. [10] J.S. Karp, S. Surti, M.E. Daube-Witherspoon, et al., J. Nucl. Med. 44 (2003) 1340. [11] C. Kuntner, E. Auffray, Chr. Dujardin, P. Lecoq, Chr. Pedrini, M. Schneegans, Development of new mixed LuYAP:ce crystals for application in a small animal PET scanner with DOI capability, IEEE Nuclear Science Symposium Conference Record, 2002; pp. M3-21. [12] C.L. Melcher, J.S. Schweitzer, IEEE Trans. Nucl. Sci. NS-39 (1992) 502. [13] S.C. Strother, M.E. Casey, E.J. Hoffman, IEEE Trans. Nucl. Sci. 37 (1990) 783. [14] D.W. Townsend, L. Byars, M. Defrise, A. Geissbuhler, R. Nutt, Phys. Med. Biol. 39 (1994) 401. [15] W.W. Moses, P.R.G. Virador, S.E. Derenzo, R.H. Huesman, T.F. Budinger, IEEE Trans. Nucl. Sci. NS-44 (1997) 1487. [16] C.S. Levin, E.J. Hoffman, Phys. Med. Biol. 44 (1999) 781; C.S. Levin, E.J. Hoffman, CORRIGENDUM Phys. Med. Biol. 45 (2000) 559.
Nuclear Instruments and Methods in Physics Research A 525 (2004) 242–248
Experience with scintillators for PET: towards the fifth generation of PET scanners L. Erikssona,b,*, D. Townsendc, M. Erikssona,b, C. Melcherd, M. Schmanda, B. Bendriema, R. Nuttd a
CPS Innovations, 810 Innovation Dr, Knoxville, TN 37932, USA b Karolinska Institute, Stockholm, Sweden c Department of Medicine and Radiology, University of Tennessee Medical Center, TN, USA d CTI Molecular Imaging, Knoxville TN, USA
Abstract Since the ECAT 2, the first commercial positron emission tomograph developed by EG&G ORTEC, four generations of scanners can be identified. The first such scanners were based on sodium iodide (NaI(Tl)) scintillators, although as early as 1978 the transition to bismuth germanate (BGO) detectors had begun. By 1981, second-generation PET scanners with up to four rings of BGO detectors were available commercially. The BGO block detector appeared in 1985, initiating the third generation of PET scanners with the potential to increase the axial coverage in a cost-effective manner. As with the second generation, the third generation of PET scanners incorporated lead septa to collimate the annihilation photons within transverse planes, thereby reducing the acquisition of scattered and random coincidences and limiting detector dead time. The fourth generation of PET scanners offered up to 15 cm axial coverage and incorporated retractable septa that permitted both 2D and 3D acquisition within the same scanner. Therefore, fifthgeneration scanner should be a fully 3D system with no septa, a 25–30 cm axial field-of-view, and a spatial resolution approaching the limits set by the physics of positron emission. Rather than using septa, limitation of randoms and scatter should be achieved directly at the detector by using a scintillator with high-light output, good energy resolution, and a fast scintillation decay time. The fast decay time will ensure low deadtime. The recent development of lutetium oxyorthosilicate (LSO), a scintillator with the required properties, suggests that a fifth generation of positron emission tomographs can now be attained. r 2004 Elsevier B.V. All rights reserved. PACS: 87.59.Vb; 87.62.+n Keywords: PET; Lutetium oxyorthosilicate; LSO; Whole-body imaging; 3D imaging; 3D data acquisition
1. Introduction
*Corresponding author. CPS Innovations, 810 Innovation Dr, Knoxville, TN 37932, USA Tel.: +1-865-218-2246; fax: +1-865-218-3000. E-mail address: [email protected] (L. Eriksson).
Positron emission tomography (PET) scanners have been available since the designs of the mid1970s that originated with the development of the PETT III [1] at Washington University in St. Louis. This initial design led to the ECAT 2
0168-9002/$ - see front matter r 2004 Elsevier B.V. All rights reserved. doi:10.1016/j.nima.2004.03.067
ARTICLE IN PRESS L. Eriksson et al. / Nuclear Instruments and Methods in Physics Research A 525 (2004) 242–248
scanner manufactured by EG&G ORTEC in Oak Ridge, Tennessee. These early designs were the forerunners of the high-performance PET technology available today. Over the past 25 years, PET scanner technology has undergone continuous and steady improvement, especially in the past 5 years since the introduction of reimbursement for PET imaging in a number of different cancers. By reviewing the progress of PET technology, four different generations can be identified even though the boundaries between consecutive generations may not always be well defined. This paper will describe the characteristics of each of the four generations and anticipate the design features of a fifth generation of PET scanner technology.
2. Four generations of pet scanners The first generation of PET scanners incorporated a single ring, or hexagonal array, of sodium iodide (NaI(Tl)) detectors. This scintillator had first been used in gamma cameras and was a natural choice for PET. However, while being an optimal scintillator at 140 keV for use in standard nuclear medicine with 99mTc, sodium iodide was less well suited to detect the higher energy, 511 keV photons from positron annihilation. A new scintillator, bismuth germanate (BGO) discovered in the early 1970s [2], had much higher sensitivity than sodium iodide at 511 keV, even though the light output was around 15% that of NaI(Tl). The potential of BGO for PET was first recognized and described in a paper by Cho et al. in 1977 [3]. The introduction of BGO gave rise to the second generation of PET scanners, initially incorporating only one or two rings of detectors. Designs with two or more detector rings, such as the Scanditronix PC-384-7b scanner, also included annular lead shields, or septa, between the detector rings. The septa served to shield the detectors from outof-plane scatter and randoms and reduce dead time by restricting coincidences to within the transaxial plane. In 1985, Casey and Nutt at CTI (Knoxville, TN) proposed the BGO detector block [4] opening up the possibility for multiple rings of detectors with small individual crystals covering an increased axial extent at a reasonable cost. The
243
block detector scanners of the mid-1980s such as the CTI 831 and the Scandotronix PC2048-15B represented the third generation of PET scanners. These designs still included septa between the detector rings to limit scatter and random coincidences and the axial coverage was either 5 cm or 10 cm depending on the number of rings of detector blocks (one or two) actually incorporated into the scanner. The fourth generation of scanners saw the introduction of retractable septa that, when fully retracted from the imaging volume, allowed coincidences to be collected between any pair of detector rings. This a 3D mode of operation increased the system sensitivity by factors of 5 or more, although scatter and randoms rates also increased. Such 3D data sets demanded a fully 3D reconstruction algorithm, the feasibility of which had earlier been demonstrated by Townsend et al. [5] by physically removing the septa from a thirdgeneration PET scanner. The first scanner with retractable septa, the ECAT 953B, primarily a brain scanner with a 35 cm patient port and 10 cm axial coverage, appeared in 1990. Results and comparisons between 2D and 3D operation of the ECAT 953B were first published in Ref [6]. By 1991, all septa-retractable fourth generation scanners incorporated an additional ring of detector blocks, thus covering 15 cm axially. Examples of fourth-generation PET scanners include the ECAT EXACT and the ECAT EXACT HR+ from CTI PET Systems (Knoxville, TN) and the ADVANCE scanner from GE Medical Systems (Milwaukee, WI). CTI PET Systems also designed the unique ECAT EXACT HR++, a PET scanner that incorporated no septa and covered an axial field of 24 cm with 48 rings of 4 mm BGO crystals. The example of this design was installed at Hammersmith Hospital, London in 1996 [7].
3. Towards a fifth generation of PET scanners A fifth generation of scanners will obviously improve and extend the performance of the fourth generation, just as the fourth generation represented an improvement over the third. Parameters to be improved include sensitivity, spatial
ARTICLE IN PRESS 244
L. Eriksson et al. / Nuclear Instruments and Methods in Physics Research A 525 (2004) 242–248
resolution, axial coverage, random rates and scatter fraction, count rate and dead time. In general terms, the next generation should have increased axial coverage of at least 30 cm to make maximum use of the emitted photon flux, operate in 3D only (no septa) for maximum sensitivity, but comprising small detector elements of 4 mm or less. High spatial resolution is essential, particularly for imaging the brain, and, as with animal PET scanners, the resolution should be increased to the limit set by the physics of positron emission. To limit the levels of random and scattered coincidences, the detector must have good coincidence timing and energy resolution, respectively. A time resolution of 1–2 ns, or better, is desirable, with an energy resolution of at least 14%. For optimal count rate performance, a low dead time is required such that the processing time for detected photons is within 100–200 ns including time and energy validation and reset time. Depending on the size of the detector units and the incoming photon flux, the processing and reset time should be reduced below 200 ns. The limitations of spatial resolution with the current block detector design are well understood. The block uses Anger logic to localize the incident photon to a particular detector element. Spatial resolution is therefore limited by the size of the detector element (pixel) in relation to the size and sampling distance of the optical sensors such as the phototubes. The maximum number of pixels per optical sensor depends strongly on the sharing of the light, the design of the light guide, and the amount of light emitted from the scintillator by conversion of a 511 keV photon. With current PET scanner technology for imaging patients, typical pixel sizes are 6 6 d mm3 down to 2 2 d mm3, where the depth d is in the range 20–30 mm. It is well known that other factors also limit spatial resolution, such as the range of the positrons after emission from the nucleus and before annihilation, and the acolinearity of the photon pairs due to the non-zero momentum of the annihilating electron–positron system. The uncertainty due to positron range is a function of the energy of the positron when ejected from the nucleus, which in turn depends upon the particular positron-emitting isotope used. The uncertainty
due to acolinearity depends on the distance between the coincident detectors. For a whole body PET scanner with a ring diameter of 80 cm, the contribution to the spatial resolution from acolinearity and positron range (for an isotope such as 18F) will total about 1.8 mm (full-width at half-maximum). To match this uncertainty due to positron physics, detector pixels should be around 2–3 mm in size. The energy resolution of the scintillator influences the fraction of scattered coincidences collected during an acquisition. Since all scintillators have a finite energy resolution, the energy of the 511 keV incident photon estimated from the total output of the optical sensors will have a range of values around the photo peak. For a 15% energy resolution, this range will be about 400–650 keV; the scanner is operated with a lower level discriminator (LLD) setting of 400 keV without loss of photo-peak sensitivity. With better energy resolution, say 12%, the LLD can be raised to 450 keV without loss of photo-peak events. Measured with a 70 cm long, 20 cm diameter polyethylene cylinder with a line source insert (NEMA NU 2-2001), the scatter fraction for an energy window of 400–650 keV is around 40%, decreasing to 30% or less with an energy window of 450– 650 keV. The goal, therefore, is to identify a scintillator with the appropriate properties to match the requirements of the fifth generation, high-performance PET scanner. The scintillator should have high light output to meet the positioning accuracy required, good energy resolution, high density and photo fraction for good sensitivity, and fast rise time and decay to ensure good coincidence timing and low detector dead time. Despite over two decades as the PET scintillator of choice, BGO is no longer a good match for these demands. The low light output (15% of NaI) and long decay time of the scintillation (300 ns) limits the timing resolution to 5–6 ns and the light signal would require integration for 500–600 ns to ensure adequate energy resolution; a long integration time results in high detector dead time. There are a number of scintillators that are possible candidates for the fifth generation of highperformance PET scanners, as summarized in
ARTICLE IN PRESS L. Eriksson et al. / Nuclear Instruments and Methods in Physics Research A 525 (2004) 242–248
Table 1 compared with BGO and NaI(Tl). One candidate, gadolinium oxyorthosilicate (Gd2SiO5:Ce, or GSO) was first discovered in 1983 [8] and used in early PET scanners in conjunction with BGO. Two layer crystals, one layer of BGO and one of GSO were positioned on the same photomultiplier and pulse shape discrimination used to separate the two scintillators [9]. Unfortunately, GSO cleaves rather easily when cut or machined so special techniques have been developed to avoid the crystals fracturing during detector construction. GSO has recently been used in a brain PET scanner, the G-PET [10], and in a commercial whole-body PET scanner, the ALLEGRO from Phillips Medical Systems. The G-PET and ALLEGRO can be considered examples of the fifth generation of PET scanners. The Crystal Clear collaboration has explored the use of cerium-activated lutetium–yttrium– ortho-aluminate (LuxY1xAlO3:Ce or LuYAP) for PET, where typically x ¼ 0:720:8 [11]. The crystal has a number of promising features such as good stopping power and a fast decay component of around 26 ns. However, for LuYAP light output is comparable to BGO, while for lutetium orthoaluminate (LuAP) it is around twice that of BGO. In either case the low light output will limit the crystal identification accuracy of the Anger logic as it does with BGO. The disadvantage of low light output is overcome by the use of the scintillator lutetium oxyorthosilicate (Lu2SiO5:Ce, or LSO), the best detector for PET known at the present time [12]. LSO combines high density and photo fraction to give good stopping power for 511 keV photons, with a light output of 75% of NaI(Tl) that is more than double that of LuAP; the
245
scintillation decay time is around 40 ns. The crystal structure is rugged and easy to machine into pixels of different dimensions, and is non-hygroscopic that facilitates packaging. LSO is an obvious choice for the fifth generation of PET scanners. The first whole-body clinical PET scanner based on LSO detectors is the ECAT ACCEL (CPS Innovations, Knoxville, TN), a design similar to the BGO-based ECAT EXACT. The relative performance of the scintillators can be assessed by comparing the count rate characteristics of scanners comprising different crystals but with similar geometric design. One metric of count rate performance is the Noise Equivalent Count Rate (NECR) [13] defined by NECR ¼
½ðT þ SÞð1 scfÞ2 T þSþR
ð1Þ
In this equation T, S and R are the true, scattered and random coincidence rates, respectively. To compare different scintillators, a fixed geometry has been used based on the ECAT EXACT design: three rings of blocks, 48 blocks per ring, a block size of 52 mm 52 mm with a depth of 25 mm; in the ECAT EXACT the BGO crystal depth is actually 20 mm. The comparison is made for the NEMA NU-2 phantom (2002), a polyethylene cylinder 20 cm in diameter and 70 cm in length with an off-center line source insert. The results of the comparison for this phantom for different scintillators are shown in Fig. 1, clearly demonstrating the superiority of LSO as the detector material for PET. The NEC curves are generated with measured sensitivity values for the different scintillators and a time window of 3 ns for LSO
Table 1 Physical properties of different scintillators for PET
Density (g/cc) Effective Z Radiation length for 511 keV (cm) Decay time (ns) Photons/Mev Light yield (% NaI) Energy resolution 662 keV in % Hygroscopic
NaI
BGO
GSO
LSO
LuYAP
3.67 50.8 2.6 230 38,000 100 7 yes
7.13 75.2 1.12 300 8200 15 10 no
6.71 59.4 1.4 30–60 10,000 25 9 no
7.4 66.4 1.14 40–43 28,000 75 9 no
7.1 61.3 1.4 20–26, 200 6000–8000 B15 8 no
ARTICLE IN PRESS 246
L. Eriksson et al. / Nuclear Instruments and Methods in Physics Research A 525 (2004) 242–248
Fig 1. NEC comparison for different scintillators with a 450 keV LLD setting based on the ECAT EXACT geometry. Only detector dead time has been modelled.
and LuYAP, 4 ns for GSO and NaI(Tl) and 11 ns for BGO. BGO has the highest sensitivity at 511 keV of all the scintillators listed in Table 1. The dead time characteristics of a BGO scanner could be substantially improved by using smaller block sizes, although at the expense of an increased number of photo detectors and more complex electronics resulting in higher production costs. The main disadvantages to continuing to use BGO for the fifth-generation PET scanners are the low light output and long scintillation time. The time resolution of a pair of detectors is given approximately by the ratio of the scintillation time to the square root of the number of photo electrons created in the photomultiplier. The best timing resolution for BGO scanners is around 5–6 ns, implying a time window of 10–12 ns to recover full sensitivity. For a fifth-generation scanner, the time window should be limited only by the size of the transverse field-of-view; a time window of 3–4 ns would cover all time-of-flight variations. The timing goal should therefore be to achieve a subnanosecond level. The next step is to identify the optimal detector configuration for the fifth-generation PET scanner. An increase in axial length could be achieved by adding extra rings of blocks to the current three or four block systems. However, while the block design has been extremely successful for almost
two decades, it may be time to consider alternatives such as multiple large flat panels that would surround the patient in a hexagonal or octagonal configuration. In such LSO-based panels, a single large (5 cm) phototube can be shared by many pixels, thus resulting in significant cost saving by reducing the overall number of phototubes. While such a design may be cost-effective, a continuous light guide (between the phototubes and the scintillator) over such a large area may lead to pixel identification errors due to pile-up effects. However, it has been found that elaborate pile-up correction techniques may be used with only a minor increase in system dead time. As a geometrical configuration, it has been shown [14] that, for a given amount of detector material, there are some advantages to arrays with large axial extent covering a limited angular range transversely as opposed to complete transverse coverage over a more limited axial extent. Obviously the advantages and disadvantages of smaller blocks as opposed to larger panels need to be carefully reviewed. An estimate for the spatial resolution (s) of such a design can be obtained from the following expression (in mm) [15]: qffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffi s ¼ 1:25 ½ðd=22 þ ð0:0022DÞ2 þ r2 þ b2 þ ðdsÞ2 ð2Þ where d is the size of the detector element (pixel), D is the separation of the detectors, r is the range of the positron and b is a measure of the crystal identification uncertainty, taken as 1 mm for LSO. A linear sampling term (ds), expressed in units of d; has also been included in Eq. (2), where s p 0.5. The results yielded by this expression for a scanner of diameter 85 cm and pixel sizes of 2, 3, and 4 mm are shown in Table 2; the positron range contribution has been estimated for imaging with isotopes 18 F and 15O. The overall sensitivity of the scanner depends not only on the physical performance of the scintillator material but also on the effective coverage of the volume around the patient into which the radiation is emitted. This coverage is defined by the detector geometry and the packing fraction, the density of pixels per detector module.
ARTICLE IN PRESS L. Eriksson et al. / Nuclear Instruments and Methods in Physics Research A 525 (2004) 242–248
The packing fraction is defined as the square of the ratio of the active crystal volume to the total detector volume. For small, 2 or 3 mm pixels, the reflective surfaces that guide the light from the scintillator into the photo sensor must be thin. The results of a calculation of the packing fraction for different pixel sizes and different reflective surfaces are shown in Fig. 2. For 2 mm pixels, a scanner resolution around 3 mm is feasible for 18F, degrading to 4 mm with 15O due to the higher energy positron. For 3 mm pixels, the estimates become 3.5 mm for 18F and 4.5 mm for 15O, values which probably represent the best spatial resoluTable 2 Estimates of reconstructed image resolution in a 85 cm fifthgeneration system as a function of detector widths and isotope used Isotope
18
15
F
O
Linear sampling factor
Detector widths (mm) 2
3
4
s¼0 s ¼ 0:25 s ¼ 0:5
3.04 3.11 3.29
3.40 3.52 3.88
3.82 4.02 4.56
s¼0 s ¼ 0:25 s ¼ 0:5
4.04 4.09 4.23
4.44 4.54 4.82
4.87 5.02 5.47
The positron ranges for 18F and 15O has been calculated using the expressions by Levin et al [16].
Fig 2. Packing fractions in coincidence for a block size of 52 52 20 mm3 block as a function of different sizes of detector elements.
247
tion that can be achieved with a whole body scanner for oncology applications. Ever since the development of the gamma camera, the principal photo detector in medical imaging has been the photomultiplier tube (PMT). For PET, the use of solid state detectors such as the avalanche photodiode (APD) has been hampered, by among other things, the low level of light emission from BGO. This situation has changed with the significantly higher light emission from LSO. PMTs, however, are faster than APDs and with the built-in amplification in the phototube itself, an inexpensive preamplifier can be used. With LSO, the ratio between the spacing and the surface area of the PMTs relative to the pixel size is in the order of 10:1 and thus 2 mm pixels can be easily identified with a 19 mm PMT. This empirical limit could be taken slightly further by the use of a slotted light guide that constrains the light into predefined regions. Since the cost of a PMT is independent of the size, large panels of 2 mm 2 mm pixels will be expensive. APDs are less bulky, but slower and require high-performance pre-amplifiers. APDs would also work in a magnetic field thus making a combined PET/MR scanner design feasible for human applications. Their quantum efficiency is around three times higher than that of phototubes and single 5 mm 5 mm APDs are available from Hamamatsu (S8664-55) with good performance characteristics. The overall package, however, is 10 mm 10 mm in size, so a high packing fraction for a large detector array would be impossible at this time. There may also be stability issues with an
Fig 3. (a) Position profile taken for one block of an HRRT panel, (b) position profile taken with four Hamamatsu APDs.
ARTICLE IN PRESS 248
L. Eriksson et al. / Nuclear Instruments and Methods in Physics Research A 525 (2004) 242–248
APD system requiring strict temperature control of the array to minimize drifts. The positioning profile for a PMT-based block is shown in Fig. 3a compared with an APD-based readout using four Hamamatsu APDs in Fig. 3b. While APDs offer considerable potential, more work is needed before the APD will routinely replace the phototube in PET scanners.
physics of positron annihilation. It remains for these, or comparable designs to bring high resolution, high-sensitivity whole-body imaging to in oncology. This is the challenge of the future in PET instrumentation.
References 4. A fifth-generation PET scanner for oncology From the above discussion it is evident that a fifth-generation whole-body PET scanner based on pixel technology and with a reconstructed spatial resolution around 3 mm is feasible. At the present time, this can only be achieved with LSO, and for cost-effectiveness, large area panels may be the solution. APDs could be an alternative to PMTs within the next 5–6 years. Finally, an example of an operational, fifth generation PET scanner is the High-Resolution Research Tomograph (HRRT), developed by CPS Innovations (Knoxville). This design comprises eight panels of 9 13 blocks viewed by a 10 14 array of 19 mm diameter PMTs; each block is 19 mm 19 mm in size and consists of two layers of different scintillators, a 10 mm layer of LSO and a 10 mm layer of LYSO, cut into 8 8 pixels per block. The layers are identified by pulse shape discrimination thus reducing the degradation of spatial resolution due to depth-of-interaction. The effective diameter of the HRRT, a scanner designed primarily for brain imaging, is around 46 cm; for a 20 cm diameter transverse field-of-view, a measured spatial resolution of 2.5 mm can be achieved with 18 F-labeled tracers. While fifth-generation PET scanner designs such as the ALLEGRO and the ACCEL are already operational in the clinical arena, it is evident that they do not attain the limits of possibility with the new scintillator materials. For specialized applications such as the brain, designs such as the G-PET and the HRRT are reaching the limits set by the
[1] M.E. Phelps, E. Hoffman, N. Mullani, C. Higgins, M. Ter-Pogossian, IEEE Trans. Biomed. Eng. NS-23 (1976) 516. [2] M.J. Weber, R.R. Monchamp, J. Appl. Phys. 44 (1973) 5495. [3] Z.H. Cho, M.R. Farukhi, J. Nucl. Med. 18 (1977) 840. [4] M. Casey, R. Nutt, IEEE Trans. Nucl. Sci. NS-33 (1986) 760. [5] D. Townsend, T. Spinks, T. Jones, A. Geissbuhler, M. Defrise, M-C. Gilardi, J. Heather, Eur. J. Nucl. Med. 15 (1989) 741. [6] D. Townsend, A. Geissbuhler, M. Defrise, E.J. Hoffman, T.J. Spinks, D.L. Bailey, M.C. Gilardi, T. Jones, IEEE Trans. Med. Imaging 10 (4) (1991) 505. [7] T. Jones, D.L. Bailey, P.M. Bloomfield, T.J. Spinks, W.F. Jones, K. Vaigneur, J. Reed, J. Young, D.F. Newport, C. Moyers, M.E. Casey, R. Nutt, J. Nucl. Med. 37 (5) (1996) 85P. [8] K. Takagi, T. Fukazawa, Appl. Phys. Lett. 42 (1983) 43. [9] S. Holte, H. Ostertag, M. Kesselberg, J. Comput. Assisted Tomogr. 11 (1987) 691. [10] J.S. Karp, S. Surti, M.E. Daube-Witherspoon, et al., J. Nucl. Med. 44 (2003) 1340. [11] C. Kuntner, E. Auffray, Chr. Dujardin, P. Lecoq, Chr. Pedrini, M. Schneegans, Development of new mixed LuYAP:ce crystals for application in a small animal PET scanner with DOI capability, IEEE Nuclear Science Symposium Conference Record, 2002; pp. M3-21. [12] C.L. Melcher, J.S. Schweitzer, IEEE Trans. Nucl. Sci. NS-39 (1992) 502. [13] S.C. Strother, M.E. Casey, E.J. Hoffman, IEEE Trans. Nucl. Sci. 37 (1990) 783. [14] D.W. Townsend, L. Byars, M. Defrise, A. Geissbuhler, R. Nutt, Phys. Med. Biol. 39 (1994) 401. [15] W.W. Moses, P.R.G. Virador, S.E. Derenzo, R.H. Huesman, T.F. Budinger, IEEE Trans. Nucl. Sci. NS-44 (1997) 1487. [16] C.S. Levin, E.J. Hoffman, Phys. Med. Biol. 44 (1999) 781; C.S. Levin, E.J. Hoffman, CORRIGENDUM Phys. Med. Biol. 45 (2000) 559.