Design considerations of nal (tl)-based pet imaging systems improving image quality and clinical versatility

Design considerations of NaI (Tl)-based PET imaging systems improving image quality and clinical versatility Y Himisch, (Received

JP Herrero

28August

1998;revised

13 September

1998;accepted25September

1998)

ne of the major concerns against the extended use of Positron Emission Tomography (PET) in clinical environment has always been its high cost, compared to that of other nuclear medicine imaging procedures. This high cost originates from both more expensive equipments and tracers, and from the fact that separate infrastructure for PET investigations (rooms, staff, equipments, etc) is often required. Therefore, PET appears to be more complicated and time consuming than other imaging techniques. As the diagnostic value of PET investigations, especially with FDG, has proven to lead to significant improvements in diagnostic accuracy in numerous areas of clinical diagnosis and in therapy control, growing interest in performing investigations in clinical environment aro’je. The industry supports this increased awareness by providing lower cost PET imaging equipments, such as rotating sector PET-scans with a reduced number of detectors [l] or dual-headed gamma camera-based coincidence systems [2]. However, the challenge of introducing these lower cost systems is to maintain the established diagnostic quality of more expensive PET systems and/or to explore the appropriate fields of use of these new systems via thorough clinical evaluations. The shifting of PET investigations towards more and more oncological applications has also resulted in the change of technical requirements associated with whole-body imaging, as the latter

0

differ from the requirements evidenced from neurological applications. It will be shown how ADAC’s NaI-based PET systems address these technical requirements by providing cost-efficient solutions for clinical environment, which challenge the image quality and efficacy of established BGO-PET systems.

TRUE DIGITAL DETECTORS IMPROVE COUNT RATE CAPABILITY One of the most significant limitation of NaI-based coincidence systems for their application in positron imaging has been their limited count rate capability [3]. Since the coincidence count rate is always only a fraction (C).5-20%) of singles rate, a tremendous enhancement in singles count rate is necessary to achieve significant improvements in coincidence performance. Table 1 illustrates the relation between both count rates for a dual-headed system, thus giving an idea of the technical challenge. The very first publications on coincidence imaging by Anger [4] or Brownell [5] already mentioned this problem. Muehllehner, Mankoff and Karp and their coworkers have especially been working on solving the issue. This has resulted in innovations such as flexible pulse shortening, local centroiding and multichannel triggering. These findings have been described in the literature over the past years [6-111. Basically, these new systems are designed to limit the event

Y HBmisch,

Tablet i. Principle

JP

Herrero

rektion ktwesn uhxidence count rate and detector

Light

singles

Scintillation

500

Time (nsec)

Fig 7. Principle of pulse shortening in 511 keV imaging (green curve and l*ectang/e) compared to 140 keV (red curve and rectangle). Due to the higher light output at 511 keV, the amount of light integrated in 200 ns at 511 keVis similar to that integrated in about 1000 ns at 140 keV.

Local

Centroid

Detector

1

--A

Detector

2

I

Fig 2. The principle of local centroiding in Nal large area detectors. T,‘lis enables the , detectors to count more than one event per time unit. RBM

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detection in both time and space. Figure 1 illustrates the principle of pulse shortening. Due to the higher light output of 511 keV photons compared to that of 140 keV photons, no information is lost due to the shortening of the integration time that has been decreased from about 1 ms to 200 ns. Note that the integration time is flexible, depending on the occurrence time of the next pulse. Figure 2 illustrates the principle of local centroiding which has led to an electronic division of the planar NaI detector into sections or centroids, enabling it to count more than one event per time unit. Of course, enhancement of counting capability requires electronics that can handle multiple events at the same time. The design of multiple trigger channels addresses this problem by providing parallel counting possibilities. These improvements have led to the design of a digital detector which is shown in figure 3. Note that each PMT is directly coupled to one A/D converter. This detector is used in both ADAC’s nuclear camera EPIC1” detectors (which are included in the MCDPET/ACrM - option of the SOLUSTM), and VERTEXTM cameras and in Curved Crystal Technology (CCT’r”) - detectors of the dedicated C-PET”” system. The principle difference in the detector design compared to the known Anger-detectors led to the improvement of singles count rate capability from about 1 million up to 2.5 million for the EPICI’” detectors, enabling true coincidence count rates of up to 15.000 cps for the MCD1’ET/AC’r>l system, one order of magnitude more than count rates that can be achieved with conventional Anger type detectors. It will be shown below that this significant improvement is not only essential for the statistical quality of the coincidence emission images but also for performing attenuation correction by measuring transmission.

INNOVATIVE CCTTM DETECTOR TECHNOLOGY FOR C-PETTM Using Nal detectors in dedicated PET systems has many

Design

advantages since today’s users are especially looking at wholebody applications in oncology, with FDG, which mainly accounts for the vast majority of PET procedures. Extremely high count rates are not such an issue; however, the ability to perform attenuation correction in wholebody studies becomes more important. With the introduction of CCTT”’ detectors on the new C-PET’” system, the quality of PET imaging with NaI has improved. The improvements associated with the use of new detectors follow: - resolution improved by about 1 mm to less than 5 mm, as compared to the hexagonal arrangement in planar detectors due to the minimized parallax effect; - no reconstruction artifact in the center of the FOV may be observed, as evidenced from multicrystal block detector systems (ie, resolution @ 0 cm and @ 1 cm are the same); - the larger ring diameter (90 cm) provides better homogeinity and a larger patient port of 56 cm; - the system has an extremely large FOV (56 cm transaxially and 25 cm axially); - the detectors form a homogenous ring around the patient with very few insensitive areas; - the full 3D operation leads to a sensitivity of very high 400 kcps/mCi/mL in the wholebody mode; - the system still benefits from the excellent energy resolution of the NaI-detectors (12%). Figure 4 shows an example of a curved crystal CCT’” - detector with the PMT’s attached. The excellent energy resolution of the detectors is shown in figure> 5. Together with an innovative energy correction algorithm and circuitry which maintains the good energy resolution up to high count rates, it is possible to use very narrow energy threshold (-15”/0 on the low energy side), which leads to a less than 30% low 3D scatter fraction (NEMA). In addition, new coincidence electronics with a time window width of only 8 ns significantly reduce the random coincidences. As an overall result of the innovations built into the C-PET”‘, the noise equivalent count rate of the

considerations

of

-~,

Nal

(T/)-based

pigital

PET

PMT BUS

IlS%EHmsrl

,OGICCamp--‘-lDl

_

Fig 3. Principle layout of the digital fPICrM-detector as and in the dedicated C-PETrb’ system.

Crystal Technology (CCT”) detector

Fig 4. Curved

LL-l

r-l

Ethernet

used in ADAC’s nuclear cameras

with PMT’s

attached.

Phobns

30000

25ODO

20000

I

0 0

100

200

300

460 Photon

SO0

Energy

6OD 700

800

000

1000

(Iv&)

Fig 5. Energy spectrum of CCJ”” -detectors showing clear separation of 57 1 keVpeak /left) and 662 keVpeak from Cs-137 !ransmission source (right), RBM

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Y HBmisch.

Detector

JP Herrero

FOV

fig 6. Geometrical arrangement of the AC’” Cs-137 attenuation correction on ,!heMCD option. a. collimated sources translating along the detector sides. b. top view of the detector surface with scanning windows for both transmission (pink) and emission contamination (blue) measurements.

system (NEC) could be improved by more than 20’S, up to 43 kcps. As shown infigure 5, the excellent energy resolution not only helps provide lower scatter fraction, it is also a prior condition to postinjection Cs-137 singles transmission, with C-PET’ and MCDPET/AC’” for the both of them, as discussed in the next section.

POST INJECTION CS-137 SINGLES TRANSMISSION IMPROVES CLINICAL IMAGE QUALITY, QUANTITATIVE ACCURACY AND ENABLES FLEXIBLE CLINICAL OPERATION to the minimal crosstalk between the 511 keV signal of the positron emitters and the 662 keV signal of the G-137 transmission source(s) in NaI-detectors as shown in figure 5, emission and transmission acquisition may take place almost simultaneoulsy. Precondition for this approach is the exceptional high count rate of the detectors, as external sources of radiation to the already present emission from the patient are added. Figure 6 a and b shows the geometrical arrangement of the Cs-137 singles tranmsmission for the MCDrET/AC”’ in a 3D view (a) and a top view of one detector (b). There are two shielded and collimated transmission sources translating on one side of each detector. The collimated fan beam is directed toward the opposing detector, scanning the volume between the detectors, while the whole arrangement rotates 360 degrees. With the full 360”-rotation and a lateral movement of the bed, truncation is avoided. Figure 6b shows a top view of the detector surface, illustrating the scanning windows for both transmission and emission contamination measurements. Note that these electronic scanning windows are only possible with a true digital detector design. The emission contamination measurement provides a correction for the crosstalk of the 511-keV emission from the patient into the 662 keV peak of the Cs-‘137 transmission sources. Due

Fig 7. Geometrical arrangement of Cs-137 singles transmission in C-PET, axial crossecBon (left) and transaxial crossection (right).

Fig 8. SinglePass’” acquisition protocols for MCD’“/AC’“’ (top) and C-PET’” (bottom). Note that the axial overlap is necessary to correct the 30 sensitivity profiles.

Design

The method is similar with CPET’” as shown in figure 7. There is one Cs-137 source rotating within the stationary detector ring, which is automatically inserted for transmission scanning and housed in a shielded container in the gantry when not in use. It rotates for 75 seconds around the patient, a 15-second emission contamination measurement in singles mode follows. Due to the shape of the collimated fan beam there is no truncation. Cs-137 post-injection singles transmission as described above enables both C-PET’” and MCD’“/ACTM to perform SinglePass’” acquisition, ie, quasi simultaneous acquisition of both emission and transmission data in one pass. Figure 8 illustrates the principle for both MCDrEr/ACr” (top) and C-PET’” (bottom). Emission and transmission are subsequently performed through AFOV. The SinglePass acquisition protocol has sthe following clinical advantages: minimized motion artifacts; no repositioning of patients; no repeated acquisition; possibility of repeating single FOV; the ability to start reconstruction already after the first FOV, as all data is acquired. The major difference between the coincidence option MCDPET/AC’” and the dedicated C-PET’ is the total acquisition time which will be about 90 minutes for a lo&cm scan length for MCDpET/AC’” and 49 minutes for C-PET’” for the same range. The time relation between emission and transmission reflects the fact that the majority of measurement time should be spent on acquiring the data, whereas corrections should only need a minimal amount of time. This relation has been distorted for a long time with conventional BGOPET-systems using Ge-68 coincidence transmission. With the SinglePass’” technology, both systems enable the user to routinely use attenuation correction for all studies. The necessity of attenuation correction in PET studies [12, 131 is still controversial. The physical effect of attenuation is not questioned; more particularly, it can be significant in the body (up to 80% attenuation). Therefore, correction of this effect, especially for quantitative or semi-quantitative

considerations

of Nal (T/J-based

PET

Uncorrected

Attenuation Corrected

/

I k.

MCvET/ACrM

3

Fig 9. Example of mediasfinal lymph node involvement in lung cancer. Top row: uncorrected emission-on/y images; bottom row: the corresponding attenuation-corrected ones (images courtesy of A Dajem, St Vincents hospital, New York, NY, USA).

Fig 10. Example of the effecl of attenuation correction in C-PETTM images: Top row: uncorrected emission-only images; bottom row: attenuation on/y; middle row: the affenuafioncorrected images.

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Y HBmisch,

JP Herrero

a

1 Fig 11. Whole-body image of a normal volunteer taken with MCDPEr/AC’” Total acquisition time (emissionttransmission) was 90 min (images courtesy of A Dayem, St Vincents hospita/, New York, NY, USA).

studies, is necessary. How even a simple diagnosis of mediastinal lymph node involvement can be affected by the presence or absence of attenuation correction is shown infiglnre 9. The top row shows the uncorrected transaxial, coronal and sagittal slices, the row below the corresponding attenuation corrected ones. Note the difference in the visibility of the mediastinal lymph nodes, especially in the transaxial slice. Total acquisition time was 90 min, usin a VERTEX”’ equiped with MCDp ET/AC”“. Figure 20 is an example of data acquired with the C-PET’“. The top row shows the uncorrected emission-only images, while the bottom row shows-the transmission-only images, and the middle row, the attenuation-corrected images. A number of artifacts can be observed in the uncorrected images, such as “hot” lungs, “cold” mediastinum, liver or skin

Fig 12. Who/e-body image of a patient with Non-Hodgkin-Lymphoma (NHL), taken with C-PET’“. mission) was 49 min (images courtesy of Reuland, EuroPET Freiburg, Germany). RRM

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Total acquisition time (emissionttrans-

Design

artifacts, and minimized uptake in the abdomen. Note how these artifacts are eliminated due to the attenuation correction [middle row). Figures 11 and 12 show two more examples of attenuation corrected images from MCDPET/AC’” @gure 11) and C-PET’” lfigure 12). Note that anatomical details are easily observed; as well, the absence of artifacts caused by attenuation may also be noted. The measurement time for the image offigure 21 was 90 min; for the C-PET’ image of figure 12, it was 49 min.

considerations

of Nal

(TI)-based

PET

they will enable more and more clinics to use the advantages of PET and a growing number of patients to benefit from the high diagnostic value of PET investigations. MCDPET/ACrM and C-PET’” offer a choice to the user to select the PET imaging option which fits best his clinical needs and the projected patient volume. Both provide cost-efficient solutions, as they minimize the cost of PET procedures by avoiding replacement of transmission sources, reducing tracer doses, and increasing the productivity of nuclear medicine departments due to shorter investiga tion times.

CONCLUSION The use of NaI-based detectors for positron imaging can lead to pertinent results in today’s clinical environment, provided that technical measures as described above are implemented. Both the coincidence option MCDPET/AC’” for SOLUS’“’ and VERTEX’” cameras as well as the dedicated C-PET’“’ system deliver useful ciinicaf imaging capabilities. Together with an adequate signal processing by Ultra SPARC”” computers and 3D iterative reconstruction i3lgOrithms, such as FORE/OSEM,

Bailey DL et al. ECAT-ART - a continuously rotating PET camera: performance characteristics, initial clinical studies and installation considerations in a nuclear medicine depariment. Ezrr ] Nut Men 1997; 14: p 6-15 Nelleman P, Hines H, Rr.jymer W, Murhllehner C, Geagan h4. Performance characteristics of a dual head SPECT scanner with PETcapability. In: 1995 IEEE iV~rcl~ar Sciellcc :;ympsiurn and Medicnl imaging Confemlcr Record; 1995 Ott 27-28; San Francisco I’helps ME. Cherry, S.R. The changing design of positron imaging; systems. Ciin Positron 7rr~agirig 7998; 1: 31-45

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Anger H. Gamma-ray and positron scintillation cameras. Nucleonics 1963; 21: 56-9 Brownell G, Burnham C. A multi-crystal positron camera. IEEE Tram Nucl Sci 1972; 19: 201-5 Muehllehner G. Positron camera with extended counting rate capability. I Nitci Med 1975; 16: 653-7 Karp JS, Muehllehner G, Beerbohm D, Mankoff DA. Event localization in a continous scintillation detector using digital processing IEEE Tram Nucl Sci 1986; 33: 550-S Karp JS, Mankoff DA, Muehllehner G. A position-sensitive detector for use in positron emission tomography. Nut/ [rlsfr M<*t/l 1988; A273: 891-7 Muehllehner G, Karp JS, Mankoff DA, Beerbohm D, Ordonez CE. Design and performance of a new positron tomograph. /EEE Tram 1998; 35: 670-4 10 Mankoff DA, Muehllehner G, Karp JS. The high count rate performance of a two-dimensionally position-sensitive detector for positron emission tomography. Pllys Med Biol; 34: 437-56 11 Mankoff DA, Muehllehner G, Miles GE. A local coincidence triggering system for PET tomographs composed of large area position-sensitive detectors. IEEE Trnns 19YO; 37: 730-6 12 Bengel et al. Whole body positron emission tomography in clinical oncology: comparison between attenuation corrected and uncorrected images. Eu, J Nucl Mcd 1997; 24: 1091-R 13 Imran et al. Lesion to brackground ratio in non-attenuation corrected whole body FDG-PET images. [ Nucl Mcd 1998; 39: 1219-23