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New high spatial resolution portable camera in medical imaging
ARTICLE IN PRESS
Nuclear Instruments and Methods in Physics Research A 577 (2007) 604–610 www.elsevier.com/locate/nima
New high spatial resolution portable camera in medical imaging C. Trottaa,b, R. Massaria,b, N. Palermob,c, F. Scopinaroa,d, A. Soluria,b, a
Institute of Biomedical Engineering, CNR, Rome, Italy b Li-tech srl, Lauzacco Pavia di Udine (UD), Italy c Microtelecom srl, Lauzacco Pavia di Udine (UD), Italy d Department of Radiological Sciences, University La Sapienza Rome, Italy Received 6 November 2006; received in revised form 14 March 2007; accepted 20 March 2007 Available online 25 April 2007
Abstract In the last years, many studies have been carried out on portable gamma cameras in order to optimize a device for medical imaging. In this paper, we present a new type of gamma camera, for low energies detection, based on a position sensitive photomultiplier tube Hamamatsu Flat Panel H8500 and an innovative technique based on CsI(Tl) scintillation crystals inserted into the square holes of a tungsten collimator. The geometrical features of this collimator–scintillator structure, which affect the camera spatial resolution and sensitivity, were chosen to offer optimal performances in clinical functional examinations. Detector sensitivity, energy resolution and spatial resolution were measured and the acquired image quality was evaluated with particular attention to the pixel identification capability. This low weight (about 2 kg) portable gamma camera was developed thanks to a miniaturized resistive chain electronic readout, combined with a dedicated compact 4 channel ADC board. This data acquisition board, designed by our research group, showed excellent performances, with respect to a commercial PCI 6110E card (National Intruments), in term of sampling period and additional on board operation for data pre-processing. r 2007 Elsevier B.V. All rights reserved. PACS: 87.58.Xs; 29.40.Mc; 85.60.Ha Keywords: Gamma ray imager; Scintillation detector; Position sensitive photomultiplier tube; Portable gamma camera
1. Introduction In the last years, new small Field of View (FOV), highresolution scintillation gamma cameras have been achieved in Nuclear Medicine thanks to the new generation of Position Sensitive Photomultiplier Tube (PSPMT). Many works, concerning this subject, were mainly oriented toward reaching the best intrinsic detector spatial and energy resolution in order to point out the performances of the PSPMT and the scintillating crystal [1]. To carry out a portable gamma camera suitable for medical application it is important to take into account other aspects, often underestimated, like the collimator, the electronic readout, Corresponding author. Institute of Biomedical Engineering— CNR, Via Salaria Km 29.300, C.P. 10 00016 Monterotondo (Roma), Italy. Tel.: +39 6 90672923; fax: +39 6 90672829. E-mail address: [email protected] (A. Soluri).
0168-9002/$ - see front matter r 2007 Elsevier B.V. All rights reserved. doi:10.1016/j.nima.2007.03.037
the data acquisition system. For example, the choice of the collimator is closely related to sensitivity and spatial resolution. In clinical application an high spatial resolution (e.g. 0.5 mm) gamma camera with a low sensitivity is suboptimal for most of medical examinations, because their duration must be as short as possible and the injected dose must be as low as possible. Another important parameter is the maximum count rate supported by the electronic readout. Actually electronic systems, nonoptimized for medical use, reduce the total number of collected events, while they increase the acquisition time. Two main methods are usually employed to readout the charge collected by PSPMT anodes. The first consists of connecting each anode of PSPMT to single, independent electronic module [2]. In this case, all the anode signals are converted and addressed to the acquisition system for data processing. The main advantages are: the possibility to correct the gain inhomogeneity of the PSPMT’s anodes and to
ARTICLE IN PRESS C. Trotta et al. / Nuclear Instruments and Methods in Physics Research A 577 (2007) 604–610
perform the study of the crystal scintillating function, whereas the disadvantages are low count rate reachable, due to multiplexing operation and the high cost related to the large number of electronic channels module. In the resistive chain method the anodes are directly connected on a proper resistive network, generally with four output channels [3]. This means reduced number of signals to amplify, sample and process, with a consequent improving in term of count rate and cost. In addition, the low density of electronic components involves a reduction of the power consuming, electronic noise, size of electronic module, and then the camera dimensions. Generally it is important to minimize the distance of the channel preamplifiers to the PSPMT’s output in order to improve the signal-to-noise ratio (SNR). Consequently, an high number of preamplifier, assembled in a miniaturized electronic module close to the PSPMT, involve high-power consuming with a consequent raising of the dark current, due to the temperature of the detector head. In this article, we present a low weight (about 2 kg), fully portable imager supplied with rechargeable batteries, suitable for medical and surgery applications. Our group developed and built many versions of new portable scintigraphic devices for medical use, patenting them since 1997 [4,5]. The proposed camera is assembled with an innovative technique based on CsI(Tl) scintillation crystals encapsulated into a square-hole collimator [6]. Here, we investigated the image performances of the Flat Panel H8500 PSPMT [7] coupled with this collimator–scintillator structure. We also developed a low-power consuming miniaturized resistive chain for the PSPMT charge read-out, and a compact dedicated ADC four channels, 12 bit, 20 Msample/s each, which allows on board data pre-processing, reducing the PC computation time. Reckoning the consumptions of the whole system and minimizing it, we were able to create the first battery supplied gamma camera with an autonomy of about 5 h. The small FOV area (about 4 in.2), the spatial resolution and the sensitivity of the proposed detector allows many clinical applications e.g.: sentinel lymph node, radioguided surgery, thyroid and para-thyroid surgical resection, diabetic foot, etc. 2. Equipment and method The camera presented in this paper is named Imaging Probe (IP). It consists of a square hole tungsten collimator integrated with a CsI(Tl) scintillator structure, described elsewhere [6], coupled with a H8500 Flat Panel photomultiplier tube, a pure tungsten shielding housing, a miniaturized charge readout electronics and a new data acquisition system for on-line imaging. The whole system is shown together with the display unit in Fig. 1. Briefly, the parallel square hole collimator is made of pure tungsten with 200 mm thick septa and it consist of a 6 mm primary block with crystals integrated and an adjunctive 24 mm block collimator arranged over the collimator–crystals integrated structure. The size of each
605
Fig. 1. The Imaging Probe and the display unit.
hole is 2.25 2.25 mm2. Building a collimator with pure tungsten implies an increase of efficiency, respect an analogous one in lead [8]. In addition tungsten septa interposed among scintillation crystals allows cross-talk suppression among them [9]. The scintillation structure is composed of 20 20 CsI(Tl) crystal array (Spectra Physics-Hilger, UK) with a FOV of 49.0 49.0 mm2. The crystals have a size of 2.05 2.05 5.0 mm3 and are covered by 100 mm of white reflective epoxy on their five blind surfaces. The space left to reach the height of collimator primary module is filled with 1 mm of white epoxy. The collimator–scintillator array was designed to reach the best trade-off between spatial resolution and sensitivity for medical use. This new building technique involves a perfect matching between collimator hole and crystal, with better SNR due to the removal of pattern mismatching. The Flat Panel PSPMT has an external size of 52.0 52.0 28.0 mm3 with an active area of 49.0 49.0 mm2, the glass windows thickness is 1.5 mm, the photocathode is bialkali and the multiplication system consist of 12 metal channel dynodes so the gain of the H8500 has a nominal value of 3 106 @1000 V, with an anode gain variation range of about 2:1. The multiplied charge is collected by an array of 8 8 anodes and the size of each anode is 6.08 6.08 mm2 except the external ones with a size of 6.28 6.28 mm2. The read-out electronic board was miniaturized (52.0 52.0 5.0 mm3) (Fig. 2 top), in order to carry out a portable camera. The same board, directly connected to the PSPMT, includes four low-noise charge preamplifier. The signals coming from the readout electronic are sampled with a new dedicated compact ADC card developed by our research group. This new data acquisition system card is connected via USB bus to a dedicated system based on embedded PC. The USB ADC-card have a sampling period of 50 ns and an acquisition time window, for each event, of
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C. Trotta et al. / Nuclear Instruments and Methods in Physics Research A 577 (2007) 604–610
Fig. 2. Resistive chain schematic diagram (top left), view of the miniaturized electronic read-out together with the PSPMT (top right) and overall sizes comparison between National 6110E and USB ADC-Card (bottom).
10 ms. During the acquisition time the USB ADC-card integrates the signals over the time. Sampling period can sample a scintillating pulse coming from a crystal faster than CsI(Tl) (e.g. NaI(Tl)). The card, developed for on board data operation, calculates the integral of each pulse signal over the acquisition time. The board is connected to the embedded PC via USB 2.0 protocol allowing a fast data transmission up to 480 Mbps (comparable with a PCI one). The embeddeb PC performs the operations to calculate the X, Y coordinates and the energy values for each event. The maximum count/rate of USB ADC-card depends on the specific scintillation crystal used and in the case of CsI(Tl) is 80,000 counts/s. We compared the USB ADC-card with a commercial PCI-6110E National Instruments board. This board was plugged into a PC (Pentium IV at 3.0 GHz) via bus PCI full size. The PCI-6110E has a sampling period of 200 ns. The time of sampling is 10 ms and all samples are transferred via bus PCI to a dedicated software (LabWindows/CVI environment) and no pre-processing operation is available. The maximum count rate measured is about 50,000 count/s. This commercial board is adequate in many standard device and we used it in previous works [6,10–13]. An acquisition system, that concern the use of the
PCI-6110E, is slowed down, with respect to the USB ADC-card, by the time necessary to the software for the data processing . The external dimensions of the USB ADC-card are roughly four times smaller than the PCI 6110E ones (Fig. 2 bottom). The software language used was C++/Linux operative system environment. It performs the corrections and calculations needed to reconstruct the scintigraphic image online, therefore the image is displayed in real time (refreshing time: 500 m/s). The software performs dynamic acquisitions. Images were reconstructed by calculating the centroid coordinates of the single event via software. All the tests were performed using a 57Co point source (1.85 MBq activity) emitting 122.1 keV (89%) and 136.5 keV (11%) gamma-ray lines and a solution of 99m Tc sodium pertechnetate (gamma-ray line at 140.5 keV). Measurements of sensitivity were performed with the 57Co point source at 5 cm distance. A Conventional Gamma Camera (CGC) was used to make a comparison with the IP. The CGC is double head gamma camera and each head is constitued by a NaI(Tl) scintillating crystal 9.5 mm thick, 55 PMT every one equipped with a dedicated ADC and it has a FOV of
ARTICLE IN PRESS C. Trotta et al. / Nuclear Instruments and Methods in Physics Research A 577 (2007) 604–610
38.1 50.8 cm2. We used this camera with a Low Energy General Purpose (LEGP) and a VerteX High Resolution (VXHR) collimators. Main characteristics of this two collimator are, respectively, hole size: 1.40 and 2.03 mm; septa thickness: 0.180 and 0.152 mm; collimator’s length 24.7 and 54.0 mm. 3. Results and discussion Fig. 3 shows the differences between two scintillation pulses at one output of the readout electronics, sampled with the PCI 6110E (200 ns sampling period) and with the USB-ADC card (50 ns sampling period). The fluctuations
Fig. 3. Scintillation pulses acquired with National 6110E (top) and with USBADC-Card (bottom).
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of both signals are not due to the electronic noise, which is visible before the rising of the signals, but to the statistical fluctuation of the emitted photons in the scintillating crystal. The smaller sampling period of the USB-ADC card allows a more accurate analysis of the shape of the scintillation pulses and a better appreciation of these fluctuations. Fig. 4a shows a 57Co raw flood field irradiation image of the IP with 20 20 scintillating elements. Only 18 18 elements are visible over the whole FOV area. Considering the integrals of the counts in the two ROIs (see Fig. 4a), we found that the integral related to the external ROI is about the double of the inner one. This means that the external ROI contains the contribution of two scintillation elements which merged themselves together into one pixel. The cause of such phenomenon is the truncation of the light distributions, coming from the external crystal, by the photocathode. It is possible to recover the external pixels with an appropriate method thus obtaining a full FOV (49.0 49.0 mm2) image as shown in Fig. 4b. This procedure is based upon the analysis of each single pixel spectrometric response. The pulse height maximum of an external crystal is located at an energy lower than an inner one [6] (Fig. 5a). So, sorting the events related to the two crystals is possible to split them, by an image shifting of the external ones, as shown in Fig. 5b. However, this method presents some faults: besides the impossibility to recover the corner crystals and the nearest ones, the pulse height separation implies some events misattribution. A gamma ray, with an energy lower than photopeak one (due for example to a Compton process inside the human body), that interact with an inner pixel crystal will be attributed to the photopeak of the related external crystal. In addition, as mentioned above, this method implies that the corner crystals and the nearest ones cannot be split because of the complete overlap of their pulse height spectra. This implies a barrel shape in the reconstructed images, as it is clearly visible in Fig. 4b.
Fig. 4. (a) Raw flood field irradiation image of the IP with 20 20 elements, (b) the same image with the external crystals splitted.
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Fig. 5. Energy spectrum of the external ROI showed in Fig. 4(a) and cross-section of two elements merged together (solid line) and splitted (dashed line) (bottom). Fig. 6. (a) Cross-section of a flood field irradiation image (18 18 elements), (b) deviation from linearity for the same row of crystals. Table 1 Measured detector and theoretical collimator spatial resolutions at different source to collimator distances Distance (cm)
Camera spatial resolution (mm)
Theoretical collimator spatial resolution (mm)
1 3 5 10
3.3270.08 4.9070.13 5.9570.15 10.9170.27
3.11 4.84 6.57 10.92
This led us to perform a further camera with 18 18 elements, removing the external crystals. A study concerning the camera overall spatial resolution was performed by placing a 57Co point source at different distances from the collimator. Table 1 summarize the results obtained with the devices at 1, 3, 5 and 10 cm source to collimator distance. As visible in the table the overall camera spatial resolution measured values are quite similar to the theoretical values obtained by the classic formula: dðl e þ bÞ le 2 le ¼ l m
Rc ¼
ð1Þ
where Rc is the collimator spatial resolution, b is the distance from the point source to the collimator, l and d are, respectively, the hole length and side, whereas le
Fig. 7. Corrected energy spectrum obtained with a irradiation of all the scintillation elements.
99m
Tc flood field
estimates the effective length due to penetration and m is the linear attenuation coefficient of the septal material. The results obtained underline that the scintillation crystals do not contribute to the broadening of the point spread function as it happens in most of scintillation cameras [14]. A cross-section of a 57Co flood field irradiation image obtained with the 18 18 camera is visibile in Fig. 6a. All the elements of the section are well identified. The peak-tovalley ratios reach very good values, with a mean of 19:1 and a maximum of 56:1. Each values was calculated as the
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Fig. 8. The Derenzo phantom imaged with the CGC (VHXR collimator) and the IP.
ratio between the peak value of an element and the mean of the two corresponding valleys. A plot of the deviation from linearity for a row of crystals belonging to the scintillation structure is shown in Fig. 6b. Each value is the difference between the crystal position evaluated from the detector and the real position of it. A non-constant behavior of this plot is due to the mismatching between the lattice of the collimator–scintillator and the anodes one. The slightly asymmetrical shape respect to the central position is related to the sub-optimal allineation between the two lattices. This excellent intrinsic spatial resolution response allowed to make a look up table in order to correct the small non linearity in the position response. After correction procedures, counting uniformity response and pulse height uniformity response resulted better than 75%. Fig. 7 shows a pulse height histogram obtained from a 99m Tc flood field irradiation of the IP 18 18 after energy calibration procedures. The measured energy resolution of this corrected spectrum is 19.1% full-width at halfmaximum (FWHM). So the energy windows for all measures was set at 710%. As expected, the proposed camera has a better photopeak sensitivity compared with that ones achievable with the CGC (710% energy window), even when equipped with a LEGP collimator. In fact the measured IP sensitivity was (211714) cps/MBq while the CGC measured (119711) cps/MBq with a LEGP collimator and (6278) cps/MBq with a VXHR collimator. To point out the resolution capability of the camera, the Derenzo Phantom, filled with a solution of 99mTc, was used and placed over the device. The diameter of the holes ranges from 2 to 4 mm while the distance between the center of two near holes is the double of the hole’s diameter (see Fig. 8). The same measure was perfomed with the CGC equipped with a VXHR collimator. As visible in Fig. 8, the IP was able to distinguish all the holes of the phantom whereas in the CGC, the holes with diameter less than 3 mm are merged together, making impossible the identification of all the objects. The better objects identification of the IP, with respect to the CGC, is not only due to a superior intrinsic spatial resolution but also to very good image contrast. The last
one parameter is a direct consequence of the innovative collimator–scintillator structure, proposed in this paper. 4. Conclusions Thanks to a miniaturized electronic readout and a compact ADC card, carried out by our research group, we developed a low weight (about 2 kg), battery supplied, really portable gamma camera with small FOV (about 4 in.2). We designed the collimator–scintillator structure trying to reach the best tradeoff between detector sensitivity and overall spatial resolution for clinical applications. We think that spatial resolutions of about 2 mm are a good compromise since under actual conditions the detectability of the lesions is always extremely difficult in the absence of specific radio-pharmaceuticals, such as to produce on the image acceptable SNR and contrast. The performances of the proposed camera are sharply better than that ones achievable with a CGC, in term of intrinsic spatial resolution, sensitivity and image contrast. The test carried out on this camera highlights that it is quite difficult to perform a large FOV (more than 4 in.2) IP assembling together arrays of Flat Panel H8500 PMTs with independent electronic modules. According to a previous work [6], a solution to this problem could be the newest R8900-00C12 Hamamatsu PSPMT with dead zones reduced down to 1.35 mm on each side. Acknowledgments We thank the technologist staff at Li-tech srl, CEA srl, Microtelecom srl, Lauzacco Pavia di Udine (UD) Italy. This work was partially done under MIUR spin-off Project ‘‘Miniaturised scintigraphic devices’’. References [1] M.N. Cinti, R. Scafe, R. Pellegrini, C. Trotta, P. Bennati, S. Ridolfi, L. Montani, F. Cusanno, F. Garibaldi, J. Telfer, R. Pani, in: IEEE Nuclear Science Symposium Conference Record 4, 2003, p. 2371 [2] F. Garibaldi, R. Accorsi, M.N. Cinti, E. Cisbani, S. Colilli, F. Cusanno, R. Fratoni, F. Giuliani, M. Gricia, R.C. Lanza, S. Lo Meo, M. Lucentini, S. Majewski, R. Pani, R. Pellegrini, F. Santavenere, B. Tsui, in: IEEE Nuclear Science Symposium Conference Record J03-5, 2005, p. 2811.
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[3] S. Siegel, R.W. Silverman, Yiping-Shao, S.R. Cherry, IEEE Trans. Nucl. Sci. NS-43 (3) (1995). [4] A. Soluri, R. Pani, Miniaturized gamma camera with very high spatial resolution, US Patent 6,242,744 B1, June 5, 2001. [5] A. Soluri, R. Pani, Flat scintillation gamma camera, with very high spatial resolution, with modular structure, US Patent 6,232,605 B1, May 15, 2001. [6] A. Soluri, R. Massari, C. Trotta, L. Montani, G. Iurlaro, A.M. Mangano, F. Scopinaro, R. Scafe`, Nucl. Instr. and Meth. A 554 (2005) 331. [7] Hamamatsu Photonics K.K., Electron Tube Center, ‘‘H8500-Data Sheet,’’ 2003. [8] G.J. Royle, N.E. Royle, R. Pani, R.D. Speller, IEEE Nucl. Sci. Symp. Conf. Reco. 3 (1995) 1584.
[9] Y. Shao, S.R. Cherry, S. Siegel, R.W. Silverman, IEEE Trans. Nucl. Sci. NS-43 (1996) 1938. [10] F. Scopinaro, R. Pani, A. Soluri, R. Pellegrini, R. Scafe‘, G. De Vincentis, F. Capoccetti, V. David, S. Chiarini, S. Stella, Tumori 86 (2000) 329. [11] A. Soluri, R. Scafe‘, F. Capoccetti, N. Burgio, A. Schiaratura, R. Pani, R. Pellegrini, M.N. Cinti, M. Mechella, A. Amanti, V. David, F. Scopinaro, Nucl. Instr. and Meth. A 497 (2003) 114. [12] A. Soluri, R. Scafe‘, F. Falcini, R. Sala, N. Burgio, G. Fiorentini, G. Giorgetti, S. Stella, S. Chiarini, F. Scopinaro, Nucl. Instr. and Meth. A 497 (2003) 122. [13] A. Soluri, F. Scopinaro, G. De Vincentis, A. Varvarigou, R. Scafe‘, R. Massa, O. Schillaci, A. Spanu, V. David, Anticancer Res. 23 (2003) 2139. [14] J.A. Sorenson, M.E. Phelps, Phys. Nucl. Med. (1980) 299.
Nuclear Instruments and Methods in Physics Research A 577 (2007) 604–610 www.elsevier.com/locate/nima
New high spatial resolution portable camera in medical imaging C. Trottaa,b, R. Massaria,b, N. Palermob,c, F. Scopinaroa,d, A. Soluria,b, a
Institute of Biomedical Engineering, CNR, Rome, Italy b Li-tech srl, Lauzacco Pavia di Udine (UD), Italy c Microtelecom srl, Lauzacco Pavia di Udine (UD), Italy d Department of Radiological Sciences, University La Sapienza Rome, Italy Received 6 November 2006; received in revised form 14 March 2007; accepted 20 March 2007 Available online 25 April 2007
Abstract In the last years, many studies have been carried out on portable gamma cameras in order to optimize a device for medical imaging. In this paper, we present a new type of gamma camera, for low energies detection, based on a position sensitive photomultiplier tube Hamamatsu Flat Panel H8500 and an innovative technique based on CsI(Tl) scintillation crystals inserted into the square holes of a tungsten collimator. The geometrical features of this collimator–scintillator structure, which affect the camera spatial resolution and sensitivity, were chosen to offer optimal performances in clinical functional examinations. Detector sensitivity, energy resolution and spatial resolution were measured and the acquired image quality was evaluated with particular attention to the pixel identification capability. This low weight (about 2 kg) portable gamma camera was developed thanks to a miniaturized resistive chain electronic readout, combined with a dedicated compact 4 channel ADC board. This data acquisition board, designed by our research group, showed excellent performances, with respect to a commercial PCI 6110E card (National Intruments), in term of sampling period and additional on board operation for data pre-processing. r 2007 Elsevier B.V. All rights reserved. PACS: 87.58.Xs; 29.40.Mc; 85.60.Ha Keywords: Gamma ray imager; Scintillation detector; Position sensitive photomultiplier tube; Portable gamma camera
1. Introduction In the last years, new small Field of View (FOV), highresolution scintillation gamma cameras have been achieved in Nuclear Medicine thanks to the new generation of Position Sensitive Photomultiplier Tube (PSPMT). Many works, concerning this subject, were mainly oriented toward reaching the best intrinsic detector spatial and energy resolution in order to point out the performances of the PSPMT and the scintillating crystal [1]. To carry out a portable gamma camera suitable for medical application it is important to take into account other aspects, often underestimated, like the collimator, the electronic readout, Corresponding author. Institute of Biomedical Engineering— CNR, Via Salaria Km 29.300, C.P. 10 00016 Monterotondo (Roma), Italy. Tel.: +39 6 90672923; fax: +39 6 90672829. E-mail address: [email protected] (A. Soluri).
0168-9002/$ - see front matter r 2007 Elsevier B.V. All rights reserved. doi:10.1016/j.nima.2007.03.037
the data acquisition system. For example, the choice of the collimator is closely related to sensitivity and spatial resolution. In clinical application an high spatial resolution (e.g. 0.5 mm) gamma camera with a low sensitivity is suboptimal for most of medical examinations, because their duration must be as short as possible and the injected dose must be as low as possible. Another important parameter is the maximum count rate supported by the electronic readout. Actually electronic systems, nonoptimized for medical use, reduce the total number of collected events, while they increase the acquisition time. Two main methods are usually employed to readout the charge collected by PSPMT anodes. The first consists of connecting each anode of PSPMT to single, independent electronic module [2]. In this case, all the anode signals are converted and addressed to the acquisition system for data processing. The main advantages are: the possibility to correct the gain inhomogeneity of the PSPMT’s anodes and to
ARTICLE IN PRESS C. Trotta et al. / Nuclear Instruments and Methods in Physics Research A 577 (2007) 604–610
perform the study of the crystal scintillating function, whereas the disadvantages are low count rate reachable, due to multiplexing operation and the high cost related to the large number of electronic channels module. In the resistive chain method the anodes are directly connected on a proper resistive network, generally with four output channels [3]. This means reduced number of signals to amplify, sample and process, with a consequent improving in term of count rate and cost. In addition, the low density of electronic components involves a reduction of the power consuming, electronic noise, size of electronic module, and then the camera dimensions. Generally it is important to minimize the distance of the channel preamplifiers to the PSPMT’s output in order to improve the signal-to-noise ratio (SNR). Consequently, an high number of preamplifier, assembled in a miniaturized electronic module close to the PSPMT, involve high-power consuming with a consequent raising of the dark current, due to the temperature of the detector head. In this article, we present a low weight (about 2 kg), fully portable imager supplied with rechargeable batteries, suitable for medical and surgery applications. Our group developed and built many versions of new portable scintigraphic devices for medical use, patenting them since 1997 [4,5]. The proposed camera is assembled with an innovative technique based on CsI(Tl) scintillation crystals encapsulated into a square-hole collimator [6]. Here, we investigated the image performances of the Flat Panel H8500 PSPMT [7] coupled with this collimator–scintillator structure. We also developed a low-power consuming miniaturized resistive chain for the PSPMT charge read-out, and a compact dedicated ADC four channels, 12 bit, 20 Msample/s each, which allows on board data pre-processing, reducing the PC computation time. Reckoning the consumptions of the whole system and minimizing it, we were able to create the first battery supplied gamma camera with an autonomy of about 5 h. The small FOV area (about 4 in.2), the spatial resolution and the sensitivity of the proposed detector allows many clinical applications e.g.: sentinel lymph node, radioguided surgery, thyroid and para-thyroid surgical resection, diabetic foot, etc. 2. Equipment and method The camera presented in this paper is named Imaging Probe (IP). It consists of a square hole tungsten collimator integrated with a CsI(Tl) scintillator structure, described elsewhere [6], coupled with a H8500 Flat Panel photomultiplier tube, a pure tungsten shielding housing, a miniaturized charge readout electronics and a new data acquisition system for on-line imaging. The whole system is shown together with the display unit in Fig. 1. Briefly, the parallel square hole collimator is made of pure tungsten with 200 mm thick septa and it consist of a 6 mm primary block with crystals integrated and an adjunctive 24 mm block collimator arranged over the collimator–crystals integrated structure. The size of each
605
Fig. 1. The Imaging Probe and the display unit.
hole is 2.25 2.25 mm2. Building a collimator with pure tungsten implies an increase of efficiency, respect an analogous one in lead [8]. In addition tungsten septa interposed among scintillation crystals allows cross-talk suppression among them [9]. The scintillation structure is composed of 20 20 CsI(Tl) crystal array (Spectra Physics-Hilger, UK) with a FOV of 49.0 49.0 mm2. The crystals have a size of 2.05 2.05 5.0 mm3 and are covered by 100 mm of white reflective epoxy on their five blind surfaces. The space left to reach the height of collimator primary module is filled with 1 mm of white epoxy. The collimator–scintillator array was designed to reach the best trade-off between spatial resolution and sensitivity for medical use. This new building technique involves a perfect matching between collimator hole and crystal, with better SNR due to the removal of pattern mismatching. The Flat Panel PSPMT has an external size of 52.0 52.0 28.0 mm3 with an active area of 49.0 49.0 mm2, the glass windows thickness is 1.5 mm, the photocathode is bialkali and the multiplication system consist of 12 metal channel dynodes so the gain of the H8500 has a nominal value of 3 106 @1000 V, with an anode gain variation range of about 2:1. The multiplied charge is collected by an array of 8 8 anodes and the size of each anode is 6.08 6.08 mm2 except the external ones with a size of 6.28 6.28 mm2. The read-out electronic board was miniaturized (52.0 52.0 5.0 mm3) (Fig. 2 top), in order to carry out a portable camera. The same board, directly connected to the PSPMT, includes four low-noise charge preamplifier. The signals coming from the readout electronic are sampled with a new dedicated compact ADC card developed by our research group. This new data acquisition system card is connected via USB bus to a dedicated system based on embedded PC. The USB ADC-card have a sampling period of 50 ns and an acquisition time window, for each event, of
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Fig. 2. Resistive chain schematic diagram (top left), view of the miniaturized electronic read-out together with the PSPMT (top right) and overall sizes comparison between National 6110E and USB ADC-Card (bottom).
10 ms. During the acquisition time the USB ADC-card integrates the signals over the time. Sampling period can sample a scintillating pulse coming from a crystal faster than CsI(Tl) (e.g. NaI(Tl)). The card, developed for on board data operation, calculates the integral of each pulse signal over the acquisition time. The board is connected to the embedded PC via USB 2.0 protocol allowing a fast data transmission up to 480 Mbps (comparable with a PCI one). The embeddeb PC performs the operations to calculate the X, Y coordinates and the energy values for each event. The maximum count/rate of USB ADC-card depends on the specific scintillation crystal used and in the case of CsI(Tl) is 80,000 counts/s. We compared the USB ADC-card with a commercial PCI-6110E National Instruments board. This board was plugged into a PC (Pentium IV at 3.0 GHz) via bus PCI full size. The PCI-6110E has a sampling period of 200 ns. The time of sampling is 10 ms and all samples are transferred via bus PCI to a dedicated software (LabWindows/CVI environment) and no pre-processing operation is available. The maximum count rate measured is about 50,000 count/s. This commercial board is adequate in many standard device and we used it in previous works [6,10–13]. An acquisition system, that concern the use of the
PCI-6110E, is slowed down, with respect to the USB ADC-card, by the time necessary to the software for the data processing . The external dimensions of the USB ADC-card are roughly four times smaller than the PCI 6110E ones (Fig. 2 bottom). The software language used was C++/Linux operative system environment. It performs the corrections and calculations needed to reconstruct the scintigraphic image online, therefore the image is displayed in real time (refreshing time: 500 m/s). The software performs dynamic acquisitions. Images were reconstructed by calculating the centroid coordinates of the single event via software. All the tests were performed using a 57Co point source (1.85 MBq activity) emitting 122.1 keV (89%) and 136.5 keV (11%) gamma-ray lines and a solution of 99m Tc sodium pertechnetate (gamma-ray line at 140.5 keV). Measurements of sensitivity were performed with the 57Co point source at 5 cm distance. A Conventional Gamma Camera (CGC) was used to make a comparison with the IP. The CGC is double head gamma camera and each head is constitued by a NaI(Tl) scintillating crystal 9.5 mm thick, 55 PMT every one equipped with a dedicated ADC and it has a FOV of
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38.1 50.8 cm2. We used this camera with a Low Energy General Purpose (LEGP) and a VerteX High Resolution (VXHR) collimators. Main characteristics of this two collimator are, respectively, hole size: 1.40 and 2.03 mm; septa thickness: 0.180 and 0.152 mm; collimator’s length 24.7 and 54.0 mm. 3. Results and discussion Fig. 3 shows the differences between two scintillation pulses at one output of the readout electronics, sampled with the PCI 6110E (200 ns sampling period) and with the USB-ADC card (50 ns sampling period). The fluctuations
Fig. 3. Scintillation pulses acquired with National 6110E (top) and with USBADC-Card (bottom).
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of both signals are not due to the electronic noise, which is visible before the rising of the signals, but to the statistical fluctuation of the emitted photons in the scintillating crystal. The smaller sampling period of the USB-ADC card allows a more accurate analysis of the shape of the scintillation pulses and a better appreciation of these fluctuations. Fig. 4a shows a 57Co raw flood field irradiation image of the IP with 20 20 scintillating elements. Only 18 18 elements are visible over the whole FOV area. Considering the integrals of the counts in the two ROIs (see Fig. 4a), we found that the integral related to the external ROI is about the double of the inner one. This means that the external ROI contains the contribution of two scintillation elements which merged themselves together into one pixel. The cause of such phenomenon is the truncation of the light distributions, coming from the external crystal, by the photocathode. It is possible to recover the external pixels with an appropriate method thus obtaining a full FOV (49.0 49.0 mm2) image as shown in Fig. 4b. This procedure is based upon the analysis of each single pixel spectrometric response. The pulse height maximum of an external crystal is located at an energy lower than an inner one [6] (Fig. 5a). So, sorting the events related to the two crystals is possible to split them, by an image shifting of the external ones, as shown in Fig. 5b. However, this method presents some faults: besides the impossibility to recover the corner crystals and the nearest ones, the pulse height separation implies some events misattribution. A gamma ray, with an energy lower than photopeak one (due for example to a Compton process inside the human body), that interact with an inner pixel crystal will be attributed to the photopeak of the related external crystal. In addition, as mentioned above, this method implies that the corner crystals and the nearest ones cannot be split because of the complete overlap of their pulse height spectra. This implies a barrel shape in the reconstructed images, as it is clearly visible in Fig. 4b.
Fig. 4. (a) Raw flood field irradiation image of the IP with 20 20 elements, (b) the same image with the external crystals splitted.
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Fig. 5. Energy spectrum of the external ROI showed in Fig. 4(a) and cross-section of two elements merged together (solid line) and splitted (dashed line) (bottom). Fig. 6. (a) Cross-section of a flood field irradiation image (18 18 elements), (b) deviation from linearity for the same row of crystals. Table 1 Measured detector and theoretical collimator spatial resolutions at different source to collimator distances Distance (cm)
Camera spatial resolution (mm)
Theoretical collimator spatial resolution (mm)
1 3 5 10
3.3270.08 4.9070.13 5.9570.15 10.9170.27
3.11 4.84 6.57 10.92
This led us to perform a further camera with 18 18 elements, removing the external crystals. A study concerning the camera overall spatial resolution was performed by placing a 57Co point source at different distances from the collimator. Table 1 summarize the results obtained with the devices at 1, 3, 5 and 10 cm source to collimator distance. As visible in the table the overall camera spatial resolution measured values are quite similar to the theoretical values obtained by the classic formula: dðl e þ bÞ le 2 le ¼ l m
Rc ¼
ð1Þ
where Rc is the collimator spatial resolution, b is the distance from the point source to the collimator, l and d are, respectively, the hole length and side, whereas le
Fig. 7. Corrected energy spectrum obtained with a irradiation of all the scintillation elements.
99m
Tc flood field
estimates the effective length due to penetration and m is the linear attenuation coefficient of the septal material. The results obtained underline that the scintillation crystals do not contribute to the broadening of the point spread function as it happens in most of scintillation cameras [14]. A cross-section of a 57Co flood field irradiation image obtained with the 18 18 camera is visibile in Fig. 6a. All the elements of the section are well identified. The peak-tovalley ratios reach very good values, with a mean of 19:1 and a maximum of 56:1. Each values was calculated as the
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Fig. 8. The Derenzo phantom imaged with the CGC (VHXR collimator) and the IP.
ratio between the peak value of an element and the mean of the two corresponding valleys. A plot of the deviation from linearity for a row of crystals belonging to the scintillation structure is shown in Fig. 6b. Each value is the difference between the crystal position evaluated from the detector and the real position of it. A non-constant behavior of this plot is due to the mismatching between the lattice of the collimator–scintillator and the anodes one. The slightly asymmetrical shape respect to the central position is related to the sub-optimal allineation between the two lattices. This excellent intrinsic spatial resolution response allowed to make a look up table in order to correct the small non linearity in the position response. After correction procedures, counting uniformity response and pulse height uniformity response resulted better than 75%. Fig. 7 shows a pulse height histogram obtained from a 99m Tc flood field irradiation of the IP 18 18 after energy calibration procedures. The measured energy resolution of this corrected spectrum is 19.1% full-width at halfmaximum (FWHM). So the energy windows for all measures was set at 710%. As expected, the proposed camera has a better photopeak sensitivity compared with that ones achievable with the CGC (710% energy window), even when equipped with a LEGP collimator. In fact the measured IP sensitivity was (211714) cps/MBq while the CGC measured (119711) cps/MBq with a LEGP collimator and (6278) cps/MBq with a VXHR collimator. To point out the resolution capability of the camera, the Derenzo Phantom, filled with a solution of 99mTc, was used and placed over the device. The diameter of the holes ranges from 2 to 4 mm while the distance between the center of two near holes is the double of the hole’s diameter (see Fig. 8). The same measure was perfomed with the CGC equipped with a VXHR collimator. As visible in Fig. 8, the IP was able to distinguish all the holes of the phantom whereas in the CGC, the holes with diameter less than 3 mm are merged together, making impossible the identification of all the objects. The better objects identification of the IP, with respect to the CGC, is not only due to a superior intrinsic spatial resolution but also to very good image contrast. The last
one parameter is a direct consequence of the innovative collimator–scintillator structure, proposed in this paper. 4. Conclusions Thanks to a miniaturized electronic readout and a compact ADC card, carried out by our research group, we developed a low weight (about 2 kg), battery supplied, really portable gamma camera with small FOV (about 4 in.2). We designed the collimator–scintillator structure trying to reach the best tradeoff between detector sensitivity and overall spatial resolution for clinical applications. We think that spatial resolutions of about 2 mm are a good compromise since under actual conditions the detectability of the lesions is always extremely difficult in the absence of specific radio-pharmaceuticals, such as to produce on the image acceptable SNR and contrast. The performances of the proposed camera are sharply better than that ones achievable with a CGC, in term of intrinsic spatial resolution, sensitivity and image contrast. The test carried out on this camera highlights that it is quite difficult to perform a large FOV (more than 4 in.2) IP assembling together arrays of Flat Panel H8500 PMTs with independent electronic modules. According to a previous work [6], a solution to this problem could be the newest R8900-00C12 Hamamatsu PSPMT with dead zones reduced down to 1.35 mm on each side. Acknowledgments We thank the technologist staff at Li-tech srl, CEA srl, Microtelecom srl, Lauzacco Pavia di Udine (UD) Italy. This work was partially done under MIUR spin-off Project ‘‘Miniaturised scintigraphic devices’’. References [1] M.N. Cinti, R. Scafe, R. Pellegrini, C. Trotta, P. Bennati, S. Ridolfi, L. Montani, F. Cusanno, F. Garibaldi, J. Telfer, R. Pani, in: IEEE Nuclear Science Symposium Conference Record 4, 2003, p. 2371 [2] F. Garibaldi, R. Accorsi, M.N. Cinti, E. Cisbani, S. Colilli, F. Cusanno, R. Fratoni, F. Giuliani, M. Gricia, R.C. Lanza, S. Lo Meo, M. Lucentini, S. Majewski, R. Pani, R. Pellegrini, F. Santavenere, B. Tsui, in: IEEE Nuclear Science Symposium Conference Record J03-5, 2005, p. 2811.
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