- Email: [email protected]
A CMOS APS for dental X-ray imaging using scintillating sensors
Nuclear Instruments and Methods in Physics Research A 460 (2001) 197–203
A CMOS APS for dental X-ray imaging using scintillating sensors . a, C.S. Peterssonb M.A. Abdallaa,b,*, C. Frojdh b
a . Mitthogskolan, Mid Sweden University, ITE, S-851 70 Sundsvall, Sweden . . Electronik, Electrum 229, S-164 40 Kista, Sweden Kungl Tekniska Hogskolan, Inst. for
Abstract In this paper we present an integrating CMOS Active Pixel Sensor (APS) circuit to be used with scintillator type X-ray sensors for intra oral dental X-ray imaging systems. Different pixel architectures were constructed to explore their performance characteristics and to study the feasibility of the development of such systems using the CMOS technology. A prototype 64 80 pixel array has been implemented in a CMOS 0.8 mm double poly n-well process with a pixel pitch of 50 mm. A spectral sensitivity measurement for the different pixels topologies, as well as measured X-ray direct absorption in the different APSs are presented. A measurement of the output signal showed a good linearity over a wide dynamic range. This chip showed that the very low sensitivity of the CMOS APSs to direct X-ray exposure adds a great advantage to the various CMOS advantages over CCD-based imaging systems. # 2001 Elsevier Science B.V. All rights reserved. Keywords: X-ray imaging; Pixel detectors; Readout electronics; CMOS APS; Radiation imaging; Photosensor; Integrating pixel
1. Introduction There are two different concepts to develop an X-ray detector with improved sensitivity for digital X-ray imaging. A silicon sensor could be used together with an efficient scintillator, or a semiconductor with high stopping for X-ray could be used [1,2]. The first digital system for intra-oral dental X-ray imaging based on a silicon CCD was presented in 1987 [3]. Despite the low absorption of X-ray in silicon, images with sufficient quality *Correspondence address. Mid Sweden University, ITE, S851 70 Sundsvall, Sweden. Tel.: +46-60-148495; fax: +46-60148456. E-mail address: [email protected] (M.A. Abdalla).
for diagnostic work were obtained at radiation doses slightly less than the doses required with film. Later systems based on CCDs coated with a scintillating layer showed increased signal-to-noise ratio (SNR) but a decrease in spatial resolution. Significant improvement in both spatial resolution and SNR of a scintillator coated CCD-based systems can be obtained by making pixel structures in the scintillating layer and optimizing other parameters of the system [4]. Despite its nearperfect performance, CCD technology has suffered from several weaknesses. One important drawback is the problems with its Charge Transfer Efficiency (CTE), which leads to readout speed and power consumption limitations. Moreover, CTE also makes CCD very sensitive to defective pixels as this can lead to an inoperative full column. This
0168-9002/01/$ - see front matter # 2001 Elsevier Science B.V. All rights reserved. PII: S 0 1 6 8 - 9 0 0 2 ( 0 0 ) 0 1 1 1 4 - 1
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can be of high importance for imaging devices where energetic particles (such as X-ray photons) are able to damage pixels. An other inherent CCD drawback is the blooming effect due to the non-perfect electrical isolation. In X-ray imaging where the sensing device is coated with a scintillator, direct absorption of X-rays, by CCDs, due to their thick sensitive depth, is a serious disadvantage that degrades the device SNR. Due to these drawbacks, an alternate approach to CCDs is to use a CMOS Active Pixel Sensor (APS) technology. The first results from a CMOS sensor for dental X-ray imaging were recently presented [5]. In this paper we present an integrating CMOS readout electronic circuit, ROIC coated with scintillating material for intra-oral dental X-ray imaging. The aim of this work is to evaluate the performance of CMOS technology as a possible replacement of the CCD systems.
2. Physical structure of the ROIC Intra oral digital dental X-ray imaging systems require somewhat special specifications. The sensor size should be at least 20 30 mm2. Due to the limited space in the mouth the dead space outside the sensitive area should be minimised. From the requirements of spatial resolution the sensor must be able to resolve objects with a size of 100 mm. In terms of Modulation Transfer Function (MTF), this means that the modulation has to be at least 10% at a spatial frequency of 10 LP/mm. Since the spacial resolution is limited by the pixel size, the conclusion is that the pixel size has to be not larger than 50 50 mm2. The electronic noise in the system should be less than the signal from one detected X-ray photon. In this prototype circuit, the active pixel array area comprises four pixel architectures as shown in Fig. 1(a) and (b). In each pixel, the transistor M1 forms the top section of a NMOS source follower stage, M2 the access transistor for reading the pixel and M3 the reset transistor, in addition to the photo-sensing element (D1 or T1), and an in-pixel capacitor C. The readout sequence is started by monitoring the
pixel array during X-ray exposure to detect the accumulation of charge. When an output is detected, the entire pixel array is reset by applying a reset signal to the reset transistor M3. The photocurrent generated by the photosensor (D1 or T1) charges the capacitor C. The voltage level on C is converted to current by the source follower M1. This current is then drawn to an output circuitry through the transistor M2 when an appropriate column address signal (Col) is applied. This simple pixel types are chosen for nondestructive readout, low read noise, compact layout and good spectral response. The inherent problem with these photodiode pixel types, namely reset noise, is overcome by an appropriate operating scheme, that uses the possibility of nondestructive read to cancel the effective reset voltage. A special layout technique, using the design of a four pixels pattern, was employed to achieve a maximum active area. In this four-pixels pattern, one pixel is flipped and then mirrored as shown in Fig. 1(c), which allows for components, transistor terminals and substrate contacts and wiring sharing; and hence reduced components and wiring. In all the pixel structures the photo-sensor element (D1 or T1) had equal sizes, and the poly/poly capacitor area was designed to share adequately reasonable areas. APS0 and APS00 are identical, using a photo-diode p-diffusion on an nwell. They differ in the reset transistor M3(using NMOS and PMOS transistors). APS1 and APS2 use an n-well on the p-substrate junction as the sensing element, but with APS2 being covered with the in-pixel poly/poly capacitor. APS3 and APS4 use photo-transistors as their sensing element. The sensing transistor base of APS4 was screened by the poly/poly capacitor.
3. Measured results and discussion The ROIC has been fabricated in a 0.8 mm double poly CMOS process, and was successfully tested. Testing the different pixels under illumination, it was found that the pixels whose photosensor was covered by the poly/poly capacitor are
M.A. Abdalla et al. / Nuclear Instruments and Methods in Physics Research A 460 (2001) 197–203
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Fig. 1. (a) Cross sections of pixels APS00, APS1 and APS3. (b) Schematics of APS0, APS00, APS1 and APS3. (c) Layout of four adjacent pixels (APS0).
blind to light; so they were excluded from further investigations.
3.1. Leakage current The output of pixels APS0, APS1 and APS3 were measured in darkness, at half-saturation for 3 s, and the results are shown in Fig. 2. It is clearly seen that the dark current effect for pixels APS0 and APS1 is negligible compared to the phototransistor pixel, APS3. The base-emitter (n-well to p-substrate) junction current, which is amplified by the transistor gain, is the main dark current contributer in the photo-transistor pixel.
The noise level superimposed over the dark output level is seen to be equal to few millivolts. However, this observed noise is mainly electronic noise including the noise from the measurement setup.
3.2. X-ray sensitivity The chip was exposed to direct X-ray exposure from a standard dental X-ray tube for 2 s to measure the X-ray direct detection. The response of the two photo-diode pixels, APS0 and APS1 is plotted in Fig. 3. It is also seen that the detection by APS0 is much lower than the detection by
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3.3. Spectral sensitivity For detection of optical signals by a sensing layer, the fraction of incident light absorbed in the layer is governed by the formula [6]: Iabs ¼ 1 expðaLÞ; Iinc where Iabs and Iinc are the absorbed and incident signals, respectively; a and L being the absorption coefficient and layer depth, respectively. Thus, for strong absorption, we have L> Fig. 2. Dark output signal of pixels APS0, APS1 and APS3 measured at half-saturation.
Fig. 3. Output signals of pixels APS0 and APS1 showing their respective X-ray direct absorption.
APS1, which is due to the difference in thickness of their sensitive layers. The measured X-ray direct detection signals from these CMOS APS structures are obviously much lower than the expected signals that are detected by CCDs from similar direct exposures. This is because of the large difference in the sensitive depth of the later compared to CMOS pixels. This low direct X-ray absorption in the CMOS pixels significantly improves the SNR, which is a great advantage of this technology over CCDs.
1 : aðhoÞ
A spectral response measurement for the different pixel architectures was performed to compare their efficiency and investigate their performance. The experiment was carried out for the light spectrum between 350 nm and 1 mm. For this CMOS process, the photo-diode in APS0 and APS00 are formed between the p-diffusion layer and the n-well (Fig. 1), whose junction depth is rather thin (0.4 mm). The above equations were evident in the measurement of their overall sensitivity, which was very low for wavelengths above 650 nm, while they showed their best response for the shorter wavelengths (in the blue light region). APS1 and APS3, on the other hand, had their detection medium being the n-well itself on the p-substrate, and the junction depth here is relatively thick (3.5 mm). APS1 and APS3 showed much higher sensitivity for the whole spectrum. The measured response is comparable to practical CCD systems. The simulated and measured relative spectral sensitivity results for APS0 and APS1 are shown in Fig. 4. The output signal from three different pixels when exposed to the same light flux (wavelength of 450 nm) was measured and plotted in Fig. 5. A good output linearity over a wide dynamic range is clearly shown. APS3 has the highest sensitivity because of its transistor gain. 3.4. Charge capacity and fill-factor To achieve a good SNR, the pixels layout was designed to accommodate a large amount of signal
M.A. Abdalla et al. / Nuclear Instruments and Methods in Physics Research A 460 (2001) 197–203
Fig. 4. Sensitivity plot for pixels APS0 and APS1.
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Fig. 6. Plot of the charge collected on the poly capacitor and diode depletion capacitance versus an applied input voltage.
Fig. 5. The output response for pixels APS0, APS1 and APS3 to light exposure of wave-length equals to 400 nm. The plot shows the output linearity over a wide dynamic range.
Fig. 7. Depletion charge on the photodiode.
charge. Using poly/poly layers, a linear capacitor size (1452 mm2) was designed giving 2.58 pF (Fig. 1c). The capacitor size is filling 58% of the pixel area, which was at the expense of the photodiode area (20 20 mm2), that contributes to 16% of the pixel area (fill factor). Except for APS3, the poly/poly capacitance value together with the photodiode capacitance was found to be too large for the accumulated charge during a standard exposure time. It was found that the photo-diode capacitance, in this particular CMOS process, is
nearly enough to store a sufficient charge. For this design it was found that the photodiode capacitance is equal to 10% of the poly capacitor, Fig. 6. The photodiode capacitance was not used because it suffers from non-linearity due to its voltage dependence. The estimated depletion charge on the photodiode depletion capacitance (APS0), and the charge collected on the poly/poly capacitor is plotted in Fig. 7. Reducing the poly/poly capacitor in relation to the photodiode area in future designs
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M.A. Abdalla et al. / Nuclear Instruments and Methods in Physics Research A 460 (2001) 197–203
Fig. 8. Gain map plot for pixel APS1 showing pixel array uniformity within 4.4%.
will increase the fill factor yielding a significant increase in the pixel efficiency. This means a reduced radiation dose can be used, in addition to improved SNR of the system. However, the exact fill factor to ploy capacitor area ratio needs further investigation with an overall system optimization. 3.5. Uniformity Non-uniformity in a pixel device can be caused either by variations in the photodetector, or by gain variations in the surrounding circuits. The output response from the different pixel array structures to a uniform flux of light was measured. The uniformity was found to be, for e.g. APS1, as shown in Fig. 8. The uniformity is within 4.4%.
4. Conclusions A CMOS APS imager prototype has been designed, fabricated and tested. The device was made to investigate the characteristics of different
pixel architectures, and to evaluate their performance as an alternative to the currently dominating CDD based systems in dental X-ray imaging. The opto-electrical performance of two photodiode-type pixels and a photo-transistor pixel was measured. The photo-transistor pixel (APS3) has the highest sensitivity to light compared to the photo-diode pixels (APS0&APS1) because of its transistor action. However, high dark current excludes it as a reliable candidate in this application. The APS pixel whose photo-sensing element is a p-diffusion/n-well (APS0) showed much lower sensitivity to light than APS1 and APS3 pixels, but it also exhibited an extremely low direct detection for X-rays making it a promising candidate for a low noise system. Our various measurements on the opto-electrical characteristics of the different pixels have set up a clear vision for an optimised final chip version. Depending on the efficiency of scintillator used in future designs, the in-pixel charge storage capacity has to be compromised with the fill factor. APS1 can be chosen as the most adequate alternative, because of its high quantum efficiency,
M.A. Abdalla et al. / Nuclear Instruments and Methods in Physics Research A 460 (2001) 197–203
linearity, low dark current, and low direct X-ray absorption. APS3, however, can be used if an additional circuitry is used to minimise its high dark current (will be investigated further in a future design). This work shows that CMOS APS technology is an excellent replacement to CCD technology in the intra-oral dental X-ray Imaging.
Acknowledgements We wish to acknowledge Per Helander and Henk Martijn at IMC-Stockholm for their valuable help. This work was performed in the
203
XIMAGE project funded by the European Commission within the BRITE-EURAM programme.
References [1] R. Irsigler et al., Nucl. Instr. and Meth. A 434 (1) (1999) 24–29. . [2] M. Abdalla, C. Frojdh, S. Petersson, Design of a CMOS readout circuit for dental X-ray imaging, ICECS’99. [3] K. Horner et al., Br Dent. J. 168 (1990) 244. . [4] C. Frojdh et al., Trans. Nucl. Sci. NS-45 (3) (1998) 374. [5] E.R. Fossum, R.H. Nixon, D. Schick, A 37 28 mm2 600k-pixel CMOS APS dental X-ray camera-on-a-chip with self-triggered readout, Presented at ISSCC, 1998. [6] J. Singh (Ed.), Optoelectronics – An Introduction to Materials and Devices, McGraw-Hill, Int. New York, 1996.
A CMOS APS for dental X-ray imaging using scintillating sensors . a, C.S. Peterssonb M.A. Abdallaa,b,*, C. Frojdh b
a . Mitthogskolan, Mid Sweden University, ITE, S-851 70 Sundsvall, Sweden . . Electronik, Electrum 229, S-164 40 Kista, Sweden Kungl Tekniska Hogskolan, Inst. for
Abstract In this paper we present an integrating CMOS Active Pixel Sensor (APS) circuit to be used with scintillator type X-ray sensors for intra oral dental X-ray imaging systems. Different pixel architectures were constructed to explore their performance characteristics and to study the feasibility of the development of such systems using the CMOS technology. A prototype 64 80 pixel array has been implemented in a CMOS 0.8 mm double poly n-well process with a pixel pitch of 50 mm. A spectral sensitivity measurement for the different pixels topologies, as well as measured X-ray direct absorption in the different APSs are presented. A measurement of the output signal showed a good linearity over a wide dynamic range. This chip showed that the very low sensitivity of the CMOS APSs to direct X-ray exposure adds a great advantage to the various CMOS advantages over CCD-based imaging systems. # 2001 Elsevier Science B.V. All rights reserved. Keywords: X-ray imaging; Pixel detectors; Readout electronics; CMOS APS; Radiation imaging; Photosensor; Integrating pixel
1. Introduction There are two different concepts to develop an X-ray detector with improved sensitivity for digital X-ray imaging. A silicon sensor could be used together with an efficient scintillator, or a semiconductor with high stopping for X-ray could be used [1,2]. The first digital system for intra-oral dental X-ray imaging based on a silicon CCD was presented in 1987 [3]. Despite the low absorption of X-ray in silicon, images with sufficient quality *Correspondence address. Mid Sweden University, ITE, S851 70 Sundsvall, Sweden. Tel.: +46-60-148495; fax: +46-60148456. E-mail address: [email protected] (M.A. Abdalla).
for diagnostic work were obtained at radiation doses slightly less than the doses required with film. Later systems based on CCDs coated with a scintillating layer showed increased signal-to-noise ratio (SNR) but a decrease in spatial resolution. Significant improvement in both spatial resolution and SNR of a scintillator coated CCD-based systems can be obtained by making pixel structures in the scintillating layer and optimizing other parameters of the system [4]. Despite its nearperfect performance, CCD technology has suffered from several weaknesses. One important drawback is the problems with its Charge Transfer Efficiency (CTE), which leads to readout speed and power consumption limitations. Moreover, CTE also makes CCD very sensitive to defective pixels as this can lead to an inoperative full column. This
0168-9002/01/$ - see front matter # 2001 Elsevier Science B.V. All rights reserved. PII: S 0 1 6 8 - 9 0 0 2 ( 0 0 ) 0 1 1 1 4 - 1
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M.A. Abdalla et al. / Nuclear Instruments and Methods in Physics Research A 460 (2001) 197–203
can be of high importance for imaging devices where energetic particles (such as X-ray photons) are able to damage pixels. An other inherent CCD drawback is the blooming effect due to the non-perfect electrical isolation. In X-ray imaging where the sensing device is coated with a scintillator, direct absorption of X-rays, by CCDs, due to their thick sensitive depth, is a serious disadvantage that degrades the device SNR. Due to these drawbacks, an alternate approach to CCDs is to use a CMOS Active Pixel Sensor (APS) technology. The first results from a CMOS sensor for dental X-ray imaging were recently presented [5]. In this paper we present an integrating CMOS readout electronic circuit, ROIC coated with scintillating material for intra-oral dental X-ray imaging. The aim of this work is to evaluate the performance of CMOS technology as a possible replacement of the CCD systems.
2. Physical structure of the ROIC Intra oral digital dental X-ray imaging systems require somewhat special specifications. The sensor size should be at least 20 30 mm2. Due to the limited space in the mouth the dead space outside the sensitive area should be minimised. From the requirements of spatial resolution the sensor must be able to resolve objects with a size of 100 mm. In terms of Modulation Transfer Function (MTF), this means that the modulation has to be at least 10% at a spatial frequency of 10 LP/mm. Since the spacial resolution is limited by the pixel size, the conclusion is that the pixel size has to be not larger than 50 50 mm2. The electronic noise in the system should be less than the signal from one detected X-ray photon. In this prototype circuit, the active pixel array area comprises four pixel architectures as shown in Fig. 1(a) and (b). In each pixel, the transistor M1 forms the top section of a NMOS source follower stage, M2 the access transistor for reading the pixel and M3 the reset transistor, in addition to the photo-sensing element (D1 or T1), and an in-pixel capacitor C. The readout sequence is started by monitoring the
pixel array during X-ray exposure to detect the accumulation of charge. When an output is detected, the entire pixel array is reset by applying a reset signal to the reset transistor M3. The photocurrent generated by the photosensor (D1 or T1) charges the capacitor C. The voltage level on C is converted to current by the source follower M1. This current is then drawn to an output circuitry through the transistor M2 when an appropriate column address signal (Col) is applied. This simple pixel types are chosen for nondestructive readout, low read noise, compact layout and good spectral response. The inherent problem with these photodiode pixel types, namely reset noise, is overcome by an appropriate operating scheme, that uses the possibility of nondestructive read to cancel the effective reset voltage. A special layout technique, using the design of a four pixels pattern, was employed to achieve a maximum active area. In this four-pixels pattern, one pixel is flipped and then mirrored as shown in Fig. 1(c), which allows for components, transistor terminals and substrate contacts and wiring sharing; and hence reduced components and wiring. In all the pixel structures the photo-sensor element (D1 or T1) had equal sizes, and the poly/poly capacitor area was designed to share adequately reasonable areas. APS0 and APS00 are identical, using a photo-diode p-diffusion on an nwell. They differ in the reset transistor M3(using NMOS and PMOS transistors). APS1 and APS2 use an n-well on the p-substrate junction as the sensing element, but with APS2 being covered with the in-pixel poly/poly capacitor. APS3 and APS4 use photo-transistors as their sensing element. The sensing transistor base of APS4 was screened by the poly/poly capacitor.
3. Measured results and discussion The ROIC has been fabricated in a 0.8 mm double poly CMOS process, and was successfully tested. Testing the different pixels under illumination, it was found that the pixels whose photosensor was covered by the poly/poly capacitor are
M.A. Abdalla et al. / Nuclear Instruments and Methods in Physics Research A 460 (2001) 197–203
199
Fig. 1. (a) Cross sections of pixels APS00, APS1 and APS3. (b) Schematics of APS0, APS00, APS1 and APS3. (c) Layout of four adjacent pixels (APS0).
blind to light; so they were excluded from further investigations.
3.1. Leakage current The output of pixels APS0, APS1 and APS3 were measured in darkness, at half-saturation for 3 s, and the results are shown in Fig. 2. It is clearly seen that the dark current effect for pixels APS0 and APS1 is negligible compared to the phototransistor pixel, APS3. The base-emitter (n-well to p-substrate) junction current, which is amplified by the transistor gain, is the main dark current contributer in the photo-transistor pixel.
The noise level superimposed over the dark output level is seen to be equal to few millivolts. However, this observed noise is mainly electronic noise including the noise from the measurement setup.
3.2. X-ray sensitivity The chip was exposed to direct X-ray exposure from a standard dental X-ray tube for 2 s to measure the X-ray direct detection. The response of the two photo-diode pixels, APS0 and APS1 is plotted in Fig. 3. It is also seen that the detection by APS0 is much lower than the detection by
200
M.A. Abdalla et al. / Nuclear Instruments and Methods in Physics Research A 460 (2001) 197–203
3.3. Spectral sensitivity For detection of optical signals by a sensing layer, the fraction of incident light absorbed in the layer is governed by the formula [6]: Iabs ¼ 1 expðaLÞ; Iinc where Iabs and Iinc are the absorbed and incident signals, respectively; a and L being the absorption coefficient and layer depth, respectively. Thus, for strong absorption, we have L> Fig. 2. Dark output signal of pixels APS0, APS1 and APS3 measured at half-saturation.
Fig. 3. Output signals of pixels APS0 and APS1 showing their respective X-ray direct absorption.
APS1, which is due to the difference in thickness of their sensitive layers. The measured X-ray direct detection signals from these CMOS APS structures are obviously much lower than the expected signals that are detected by CCDs from similar direct exposures. This is because of the large difference in the sensitive depth of the later compared to CMOS pixels. This low direct X-ray absorption in the CMOS pixels significantly improves the SNR, which is a great advantage of this technology over CCDs.
1 : aðhoÞ
A spectral response measurement for the different pixel architectures was performed to compare their efficiency and investigate their performance. The experiment was carried out for the light spectrum between 350 nm and 1 mm. For this CMOS process, the photo-diode in APS0 and APS00 are formed between the p-diffusion layer and the n-well (Fig. 1), whose junction depth is rather thin (0.4 mm). The above equations were evident in the measurement of their overall sensitivity, which was very low for wavelengths above 650 nm, while they showed their best response for the shorter wavelengths (in the blue light region). APS1 and APS3, on the other hand, had their detection medium being the n-well itself on the p-substrate, and the junction depth here is relatively thick (3.5 mm). APS1 and APS3 showed much higher sensitivity for the whole spectrum. The measured response is comparable to practical CCD systems. The simulated and measured relative spectral sensitivity results for APS0 and APS1 are shown in Fig. 4. The output signal from three different pixels when exposed to the same light flux (wavelength of 450 nm) was measured and plotted in Fig. 5. A good output linearity over a wide dynamic range is clearly shown. APS3 has the highest sensitivity because of its transistor gain. 3.4. Charge capacity and fill-factor To achieve a good SNR, the pixels layout was designed to accommodate a large amount of signal
M.A. Abdalla et al. / Nuclear Instruments and Methods in Physics Research A 460 (2001) 197–203
Fig. 4. Sensitivity plot for pixels APS0 and APS1.
201
Fig. 6. Plot of the charge collected on the poly capacitor and diode depletion capacitance versus an applied input voltage.
Fig. 5. The output response for pixels APS0, APS1 and APS3 to light exposure of wave-length equals to 400 nm. The plot shows the output linearity over a wide dynamic range.
Fig. 7. Depletion charge on the photodiode.
charge. Using poly/poly layers, a linear capacitor size (1452 mm2) was designed giving 2.58 pF (Fig. 1c). The capacitor size is filling 58% of the pixel area, which was at the expense of the photodiode area (20 20 mm2), that contributes to 16% of the pixel area (fill factor). Except for APS3, the poly/poly capacitance value together with the photodiode capacitance was found to be too large for the accumulated charge during a standard exposure time. It was found that the photo-diode capacitance, in this particular CMOS process, is
nearly enough to store a sufficient charge. For this design it was found that the photodiode capacitance is equal to 10% of the poly capacitor, Fig. 6. The photodiode capacitance was not used because it suffers from non-linearity due to its voltage dependence. The estimated depletion charge on the photodiode depletion capacitance (APS0), and the charge collected on the poly/poly capacitor is plotted in Fig. 7. Reducing the poly/poly capacitor in relation to the photodiode area in future designs
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Fig. 8. Gain map plot for pixel APS1 showing pixel array uniformity within 4.4%.
will increase the fill factor yielding a significant increase in the pixel efficiency. This means a reduced radiation dose can be used, in addition to improved SNR of the system. However, the exact fill factor to ploy capacitor area ratio needs further investigation with an overall system optimization. 3.5. Uniformity Non-uniformity in a pixel device can be caused either by variations in the photodetector, or by gain variations in the surrounding circuits. The output response from the different pixel array structures to a uniform flux of light was measured. The uniformity was found to be, for e.g. APS1, as shown in Fig. 8. The uniformity is within 4.4%.
4. Conclusions A CMOS APS imager prototype has been designed, fabricated and tested. The device was made to investigate the characteristics of different
pixel architectures, and to evaluate their performance as an alternative to the currently dominating CDD based systems in dental X-ray imaging. The opto-electrical performance of two photodiode-type pixels and a photo-transistor pixel was measured. The photo-transistor pixel (APS3) has the highest sensitivity to light compared to the photo-diode pixels (APS0&APS1) because of its transistor action. However, high dark current excludes it as a reliable candidate in this application. The APS pixel whose photo-sensing element is a p-diffusion/n-well (APS0) showed much lower sensitivity to light than APS1 and APS3 pixels, but it also exhibited an extremely low direct detection for X-rays making it a promising candidate for a low noise system. Our various measurements on the opto-electrical characteristics of the different pixels have set up a clear vision for an optimised final chip version. Depending on the efficiency of scintillator used in future designs, the in-pixel charge storage capacity has to be compromised with the fill factor. APS1 can be chosen as the most adequate alternative, because of its high quantum efficiency,
M.A. Abdalla et al. / Nuclear Instruments and Methods in Physics Research A 460 (2001) 197–203
linearity, low dark current, and low direct X-ray absorption. APS3, however, can be used if an additional circuitry is used to minimise its high dark current (will be investigated further in a future design). This work shows that CMOS APS technology is an excellent replacement to CCD technology in the intra-oral dental X-ray Imaging.
Acknowledgements We wish to acknowledge Per Helander and Henk Martijn at IMC-Stockholm for their valuable help. This work was performed in the
203
XIMAGE project funded by the European Commission within the BRITE-EURAM programme.
References [1] R. Irsigler et al., Nucl. Instr. and Meth. A 434 (1) (1999) 24–29. . [2] M. Abdalla, C. Frojdh, S. Petersson, Design of a CMOS readout circuit for dental X-ray imaging, ICECS’99. [3] K. Horner et al., Br Dent. J. 168 (1990) 244. . [4] C. Frojdh et al., Trans. Nucl. Sci. NS-45 (3) (1998) 374. [5] E.R. Fossum, R.H. Nixon, D. Schick, A 37 28 mm2 600k-pixel CMOS APS dental X-ray camera-on-a-chip with self-triggered readout, Presented at ISSCC, 1998. [6] J. Singh (Ed.), Optoelectronics – An Introduction to Materials and Devices, McGraw-Hill, Int. New York, 1996.