CA2295893A1 - Method for producing images in digital dental radiography - Google Patents
Method for producing images in digital dental radiography Download PDFInfo
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- CA2295893A1 CA2295893A1 CA002295893A CA2295893A CA2295893A1 CA 2295893 A1 CA2295893 A1 CA 2295893A1 CA 002295893 A CA002295893 A CA 002295893A CA 2295893 A CA2295893 A CA 2295893A CA 2295893 A1 CA2295893 A1 CA 2295893A1
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Abstract
The inventive device for digital dental radiography has a source of X-radiation, a detector device, a correcting device and a temperature control device. Said detector device consists of a CCD array on which a scintillator layer is arranged and the radiation geometry of the detector device is constant in relation to the source of X-radiation. The correcting device corrects the electrical signals generated by the CCD array by compensating fluctuations in the electrical signals of the individual elements of the CCD
array caused by a dark current of said array elements, by the array elements having different degrees of conversion efficiency and by the inhomogeneity of the scintillator layer. The temperature control device ensures that the detector device remains at a constant temperature during the detection of reference signals and image signals.
array caused by a dark current of said array elements, by the array elements having different degrees of conversion efficiency and by the inhomogeneity of the scintillator layer. The temperature control device ensures that the detector device remains at a constant temperature during the detection of reference signals and image signals.
Description
Method for Producing Images in Digital Dental Radiography Description The present invention relates to a method for producing images in digital dental radiography, and in particular to a method wherein a CCD detector device is used to detect the X-radiation.
Dental X-ray diagnostic units for which image capture is achieved digitally using CCD detectors (CCD = charge coupled device) have been known for a long time. By applying suitable clock signals, the image information is read out of the CCD detector, is preprocessed and digitalized and is then transmitted to a computer system, e.g. a personal computer, to be displayed and stored. The advantage of this method over traditional film technology resides principally in the much faster image capture.
A disadvantage of the known method, however, is that the image impressed on the detector is not reproduced exactly but is distorted by the so-called "fixed pattern noise".
This means that, for the same radiation intensity, the digital grey value of an image element, i.e. of a pixel, varies, sometimes considerably, from pixel to pixel due to manufacture-related differences. As the term "fixed pattern noise" suggests, for somebody viewing the image the manufacture-related differences manifest themselves in part as an additional noise component even though they are not of a stochastic nature. The causes of the fixed pattern noise are to be found in part in the fact that the individual pixels or elements of a CCD detector array exhibit a manufacture-related spread both in the dark signal and in the efficiency of the light conversion, i.e. in the transformation of the incident light into an electrical signal.
In most cases the incident X-radiation will not be converted directly into an electrical signal by the CCD converter.
Rather, the upper side of the detector is provided with a scintillator layer which converts incident X-radiation into visible light, which in turn is converted into an electrical signal by the CCD detector.
As a result of inhomogeneities in the scintillator layer, the image does not exhibit a uniform grey value under homogeneous irradiation but seems "patchy". In the X-ray pictures of teeth these "patches" can negatively influence the diagnosis since the dentist may interpret them falsely, e.g. as caries.
To minimize the effect of these manufacture-related differences in the dark signal and in the efficiency of the light conversion between the pixels of a detector array as well as the effect of scintillator layer inhomogeneities, relatively high X-ray doses are used in known dental X-ray diagnostic units. The effect on the created image of a temperature dependence of the dark signal and of the efficiency of the light conversion of the individual pixels is also reduced thereby. It is evidently desirable to use lower radiation doses in dental X-ray diagnosis in order to reduce the stress on the patients' health.
In the Patent Abstracts of Japan, Vol. 95, Nr. 006, 31 July 1995 and the related JP 07 072256 A a circuit for filtering out the image noise caused by capacitive errors of the read amplifiers of a semiconductor detector array for detecting X-radiation is described. Here respective signals are first detected when the respective detector element is not exposed to X-radiation or is exposed to a constant X-radiation, these signals being stored in a correction table.
Subsequently an offset and gain correction for detected image signals is performed on the basis of the correction table.
The Patent Abstracts of Japan, Vol. 16, Nr. 357 (P-1395), 31 July 1992 and the related JP 04 110690 A disclose a temperature control for a semiconductor detector array for detecting X-radiation wherein the temperature of the detection range of the semiconductor detector array is so controlled as to remain constant.
Starting from the cited prior art it is the object of the present invention to provide a method for producing images in digital dental radiography for which reductions in the quality of the created image, which are caused by manufacture-related differences in the dark signal and in the efficiency of the light conversion of the individual detector elements and by scintillator layer inhomogeneity, are eliminated.
This object is achieved by a method according to claim 1.
The method according to the present invention is based on the use of a pixelwise correction of the electrical signals generated by each element of a detector device so as to optimize the image quality, compensation of fluctations in detected image signals caused by the inhomogeneity of the scintillator layer being part of this process.
To achieve this compensation, reference signals corresponding to the output signals of the CCD array are detected when the detector device is either not exposed to X-radiation or is exposed to homogeneous X-radiation. On the basis of these reference signals the electrical signals corresponding to an image of an object are corrected. During the detection of the reference signals and the signals corresponding to the image of an object the detector device is kept at a constant temperature. In this way the effect on the resulting image of the temperature dependence of the dark current or of the conversion efficiency of the individual pixel elements can be eliminated.
The present invention thus provides a method for producing images in digital dental radiography with which, despite a reduced x-ray dose, high quality images can be generated in which a deterioration in the image quality caused by manufacture-related differences in the dark current and in the conversion efficiency of the individual image elements of the detector device and also by the inhomogeneity of the scintillator layer is prevented.
In the following an embodiment of the present invention will be described in more detail.
As has been described above, manufacture-related differences affect both the dark signal of an image element, i.e. the grey value of the pixel when not irradiated, and the conversion efficiency at a specified irradiation, i.e. the grey value of the pixel at the specified irradiation. The electrical signal generated by CCD detectors is substantially linearly dependent on the illumination intensity. The digital grey value Gwij of a pixel in a row i and a column j of a detector array can thus be described by the following equation at a specified radiation intensity:
Gwij(I)=Oij+GijI (1) where Gwij is the grey value, Oij is the offset which is caused by the dark current, Gij is the gain, and I is the intensity of the radiation. Oij and Gij can vary from pixel to pixel.
- 4a -A further problem arises from the fact that the offset Oil is strongly temperature dependent, so that the grey value of an image element depends not only on the radiation intensity but also on the temperature:
Gwi~(T,I)=Oi~(T)+Gi~I
To correct the fixed pattern noise for a specified temperature T, the values of Oil and Gig are determined for each pixel, i.e. for each detector element of the detector array. To determine these, two images are needed: an image without radiation and an image obtained with a known radiation. The known radiation may e.g. be the maximum permissible radiation.
These images, in the absence of radiation and with a known radiation, which are used to suppress the quanta-related image noise are advantageously generated by averaging several exposures under constant conditions. From the preferably averaged images the grey values without radiation and with known radiation I* are known for each image element:
Gwi~(T,I=0)=Oi~(T) ( Gwi~(T,I=I*)=Oi~(T)+Gi~I* (4) Gwi~(T,I=0) is the grey value of the pixel in row i and column j of the detector when the detector device is not exposed to X-radiation by a source of X-radiation. This grey value corresponds to the temperature-dependent offset Oi~(T) of the pixel. Gwi~(T,I=I*) is the grey value of the pixel when exposed to X-radiation with intensity I*. This grey value is composed of the offset Oil and the intensity I*
multiplied by the gain Gig of the pixel.
By subtracting equation (3) from equation (4) pixel by pixel the grey value of each pixel is obtained without the offset:
Gw~i~(T,I=I*)=Gwi~(T,I=I*)-Gwi~(T,I=0)=Gi~I* (5) To generate a high quality image by means of a detector array, all the pixels, i.e. detector elements, of the array must have the same grey value Gwmax when they are irradiated with the same maximum dose. Differences in sensitivity can be eliminated by normalizing the values Gij. From the equation Gwmax-Gw.ij(T,I=I*)Gij' (6) the normalized gain Gij' can be calculated for each pixel.
Gwmax constitutes a nominal value at maximum irradiation which all the pixels of the detector array should exhibit.
The correction value Gij' thus specifies a factor with which the grey value of each pixel must be multiplied in order that the same grey value is output for each pixel under uniform irradiation of all the detector elements.
The correction of an image characterized by the grey values Gwij occurs as follows: the corrected grey values Gw.ij result from the pixelwise subtraction of the offset values Oij of the pixels from the grey pixel values representing an image and subsequent pixelwise multiplication with the gain normalization values Gij'.
Gw'ij=(Gwij-~ij)Gij~ ( The equations (1) to (7) form the basis for improving the image quality for the device and the method according to the present invention. These equations are used according to the present invention for the pixelwise correction of the gain and the offset in the field of dental digital radiography.
As has already been described above, the offset values Oij and thus also the gain normalization values Gij' are strongly dependent on temperature. If the offset values and the gain normalization values are determined at a particular temperature, but the image for correction is recorded at some other temperature, the image quality will not be improved but will almost certainly be degraded. For this reason the temperature of the CCD detector must be held constant during the detection of the reference signals and the detection of the signals representing an image. A
constant detector temperature can be achieved in different ways.
The rear side of the CCD detector can be provided with a heating element, the temperature of the detector being held constant through active regulation. To measure the temperature, temperature detectors e.g. can be additionally arranged on the carrier of the CCD detector or be integrated into the CCD detector. If unilluminated CCD elements, so-called "dark reference pixels", are arranged on the detector, these elements can be utilized for temperature measurement because of the strong temperature dependence of the dark current.
Furthermore, the temperature can, where appropriate, be controlled by controlling the clock frequency of the CCD
detector. In this case too the temperature of the detector device can be detected, e.g. by means of temperature detectors additionally arranged on the carrier of the CCD
detector, so as to enable regulated control of the temperature of the detector device.
In the case of intraoral X-ray diagnostic units, the detector element can be held at a predetermined temperature, e.g. in a water bath, between two X-ray pictures. This predetermined temperature may e.g. be the body temperature.
In this way it is ensured that the detector element does not cool down between the X-ray pictures.
A further important condition for the successful application of the gain/offset correction is that the relative position of the detector device in relation to the source of X-radiation remains constant, or at least that it is guaranteed that the X-ray dose power is constant over the whole detector area in order to avoid a decrease in the _ $ _ image brightness towards the edge of the image, a so-called "shading". If the radiation geometry of the detector device is not constant in relation to the source of x-radiation, the gain correction will not be performed correctly, resulting in the creation of artificial shading. To avoid this problem, an arrangement analogous to a film holder in which the central beam of the source of X-radiation strikes the middle of the CCD detector can e.g. be employed.
The present invention thus provides an optimization of the image quality in the field of digital dental radiography through the use of a pixelwise gain and offset correction at a constant temperature of the detector device. The pixelwise gain/offset correction eliminates the consequences of manufacture-related differences in the dark current and in the conversion efficiency of the individual image elements and also the consequences of inhomogeneities in the scintillator layer arranged on the CCD detector.
Prerequisites for the improvement of the image quality through a pixelwise gain/offset correction are a) a constant temperature of the CCD detector and b) a constant radiation geometry or a homogeneous irradiation. The present invention thus enables the generation of high quality images at radiation intensities which are small compared with the x-ray doses normally used.
Instead of the CCD array described here, a photodiode array or a charge injection device or a CMOS image detector array can also be used.
Dental X-ray diagnostic units for which image capture is achieved digitally using CCD detectors (CCD = charge coupled device) have been known for a long time. By applying suitable clock signals, the image information is read out of the CCD detector, is preprocessed and digitalized and is then transmitted to a computer system, e.g. a personal computer, to be displayed and stored. The advantage of this method over traditional film technology resides principally in the much faster image capture.
A disadvantage of the known method, however, is that the image impressed on the detector is not reproduced exactly but is distorted by the so-called "fixed pattern noise".
This means that, for the same radiation intensity, the digital grey value of an image element, i.e. of a pixel, varies, sometimes considerably, from pixel to pixel due to manufacture-related differences. As the term "fixed pattern noise" suggests, for somebody viewing the image the manufacture-related differences manifest themselves in part as an additional noise component even though they are not of a stochastic nature. The causes of the fixed pattern noise are to be found in part in the fact that the individual pixels or elements of a CCD detector array exhibit a manufacture-related spread both in the dark signal and in the efficiency of the light conversion, i.e. in the transformation of the incident light into an electrical signal.
In most cases the incident X-radiation will not be converted directly into an electrical signal by the CCD converter.
Rather, the upper side of the detector is provided with a scintillator layer which converts incident X-radiation into visible light, which in turn is converted into an electrical signal by the CCD detector.
As a result of inhomogeneities in the scintillator layer, the image does not exhibit a uniform grey value under homogeneous irradiation but seems "patchy". In the X-ray pictures of teeth these "patches" can negatively influence the diagnosis since the dentist may interpret them falsely, e.g. as caries.
To minimize the effect of these manufacture-related differences in the dark signal and in the efficiency of the light conversion between the pixels of a detector array as well as the effect of scintillator layer inhomogeneities, relatively high X-ray doses are used in known dental X-ray diagnostic units. The effect on the created image of a temperature dependence of the dark signal and of the efficiency of the light conversion of the individual pixels is also reduced thereby. It is evidently desirable to use lower radiation doses in dental X-ray diagnosis in order to reduce the stress on the patients' health.
In the Patent Abstracts of Japan, Vol. 95, Nr. 006, 31 July 1995 and the related JP 07 072256 A a circuit for filtering out the image noise caused by capacitive errors of the read amplifiers of a semiconductor detector array for detecting X-radiation is described. Here respective signals are first detected when the respective detector element is not exposed to X-radiation or is exposed to a constant X-radiation, these signals being stored in a correction table.
Subsequently an offset and gain correction for detected image signals is performed on the basis of the correction table.
The Patent Abstracts of Japan, Vol. 16, Nr. 357 (P-1395), 31 July 1992 and the related JP 04 110690 A disclose a temperature control for a semiconductor detector array for detecting X-radiation wherein the temperature of the detection range of the semiconductor detector array is so controlled as to remain constant.
Starting from the cited prior art it is the object of the present invention to provide a method for producing images in digital dental radiography for which reductions in the quality of the created image, which are caused by manufacture-related differences in the dark signal and in the efficiency of the light conversion of the individual detector elements and by scintillator layer inhomogeneity, are eliminated.
This object is achieved by a method according to claim 1.
The method according to the present invention is based on the use of a pixelwise correction of the electrical signals generated by each element of a detector device so as to optimize the image quality, compensation of fluctations in detected image signals caused by the inhomogeneity of the scintillator layer being part of this process.
To achieve this compensation, reference signals corresponding to the output signals of the CCD array are detected when the detector device is either not exposed to X-radiation or is exposed to homogeneous X-radiation. On the basis of these reference signals the electrical signals corresponding to an image of an object are corrected. During the detection of the reference signals and the signals corresponding to the image of an object the detector device is kept at a constant temperature. In this way the effect on the resulting image of the temperature dependence of the dark current or of the conversion efficiency of the individual pixel elements can be eliminated.
The present invention thus provides a method for producing images in digital dental radiography with which, despite a reduced x-ray dose, high quality images can be generated in which a deterioration in the image quality caused by manufacture-related differences in the dark current and in the conversion efficiency of the individual image elements of the detector device and also by the inhomogeneity of the scintillator layer is prevented.
In the following an embodiment of the present invention will be described in more detail.
As has been described above, manufacture-related differences affect both the dark signal of an image element, i.e. the grey value of the pixel when not irradiated, and the conversion efficiency at a specified irradiation, i.e. the grey value of the pixel at the specified irradiation. The electrical signal generated by CCD detectors is substantially linearly dependent on the illumination intensity. The digital grey value Gwij of a pixel in a row i and a column j of a detector array can thus be described by the following equation at a specified radiation intensity:
Gwij(I)=Oij+GijI (1) where Gwij is the grey value, Oij is the offset which is caused by the dark current, Gij is the gain, and I is the intensity of the radiation. Oij and Gij can vary from pixel to pixel.
- 4a -A further problem arises from the fact that the offset Oil is strongly temperature dependent, so that the grey value of an image element depends not only on the radiation intensity but also on the temperature:
Gwi~(T,I)=Oi~(T)+Gi~I
To correct the fixed pattern noise for a specified temperature T, the values of Oil and Gig are determined for each pixel, i.e. for each detector element of the detector array. To determine these, two images are needed: an image without radiation and an image obtained with a known radiation. The known radiation may e.g. be the maximum permissible radiation.
These images, in the absence of radiation and with a known radiation, which are used to suppress the quanta-related image noise are advantageously generated by averaging several exposures under constant conditions. From the preferably averaged images the grey values without radiation and with known radiation I* are known for each image element:
Gwi~(T,I=0)=Oi~(T) ( Gwi~(T,I=I*)=Oi~(T)+Gi~I* (4) Gwi~(T,I=0) is the grey value of the pixel in row i and column j of the detector when the detector device is not exposed to X-radiation by a source of X-radiation. This grey value corresponds to the temperature-dependent offset Oi~(T) of the pixel. Gwi~(T,I=I*) is the grey value of the pixel when exposed to X-radiation with intensity I*. This grey value is composed of the offset Oil and the intensity I*
multiplied by the gain Gig of the pixel.
By subtracting equation (3) from equation (4) pixel by pixel the grey value of each pixel is obtained without the offset:
Gw~i~(T,I=I*)=Gwi~(T,I=I*)-Gwi~(T,I=0)=Gi~I* (5) To generate a high quality image by means of a detector array, all the pixels, i.e. detector elements, of the array must have the same grey value Gwmax when they are irradiated with the same maximum dose. Differences in sensitivity can be eliminated by normalizing the values Gij. From the equation Gwmax-Gw.ij(T,I=I*)Gij' (6) the normalized gain Gij' can be calculated for each pixel.
Gwmax constitutes a nominal value at maximum irradiation which all the pixels of the detector array should exhibit.
The correction value Gij' thus specifies a factor with which the grey value of each pixel must be multiplied in order that the same grey value is output for each pixel under uniform irradiation of all the detector elements.
The correction of an image characterized by the grey values Gwij occurs as follows: the corrected grey values Gw.ij result from the pixelwise subtraction of the offset values Oij of the pixels from the grey pixel values representing an image and subsequent pixelwise multiplication with the gain normalization values Gij'.
Gw'ij=(Gwij-~ij)Gij~ ( The equations (1) to (7) form the basis for improving the image quality for the device and the method according to the present invention. These equations are used according to the present invention for the pixelwise correction of the gain and the offset in the field of dental digital radiography.
As has already been described above, the offset values Oij and thus also the gain normalization values Gij' are strongly dependent on temperature. If the offset values and the gain normalization values are determined at a particular temperature, but the image for correction is recorded at some other temperature, the image quality will not be improved but will almost certainly be degraded. For this reason the temperature of the CCD detector must be held constant during the detection of the reference signals and the detection of the signals representing an image. A
constant detector temperature can be achieved in different ways.
The rear side of the CCD detector can be provided with a heating element, the temperature of the detector being held constant through active regulation. To measure the temperature, temperature detectors e.g. can be additionally arranged on the carrier of the CCD detector or be integrated into the CCD detector. If unilluminated CCD elements, so-called "dark reference pixels", are arranged on the detector, these elements can be utilized for temperature measurement because of the strong temperature dependence of the dark current.
Furthermore, the temperature can, where appropriate, be controlled by controlling the clock frequency of the CCD
detector. In this case too the temperature of the detector device can be detected, e.g. by means of temperature detectors additionally arranged on the carrier of the CCD
detector, so as to enable regulated control of the temperature of the detector device.
In the case of intraoral X-ray diagnostic units, the detector element can be held at a predetermined temperature, e.g. in a water bath, between two X-ray pictures. This predetermined temperature may e.g. be the body temperature.
In this way it is ensured that the detector element does not cool down between the X-ray pictures.
A further important condition for the successful application of the gain/offset correction is that the relative position of the detector device in relation to the source of X-radiation remains constant, or at least that it is guaranteed that the X-ray dose power is constant over the whole detector area in order to avoid a decrease in the _ $ _ image brightness towards the edge of the image, a so-called "shading". If the radiation geometry of the detector device is not constant in relation to the source of x-radiation, the gain correction will not be performed correctly, resulting in the creation of artificial shading. To avoid this problem, an arrangement analogous to a film holder in which the central beam of the source of X-radiation strikes the middle of the CCD detector can e.g. be employed.
The present invention thus provides an optimization of the image quality in the field of digital dental radiography through the use of a pixelwise gain and offset correction at a constant temperature of the detector device. The pixelwise gain/offset correction eliminates the consequences of manufacture-related differences in the dark current and in the conversion efficiency of the individual image elements and also the consequences of inhomogeneities in the scintillator layer arranged on the CCD detector.
Prerequisites for the improvement of the image quality through a pixelwise gain/offset correction are a) a constant temperature of the CCD detector and b) a constant radiation geometry or a homogeneous irradiation. The present invention thus enables the generation of high quality images at radiation intensities which are small compared with the x-ray doses normally used.
Instead of the CCD array described here, a photodiode array or a charge injection device or a CMOS image detector array can also be used.
Claims (8)
1. A method for producing images in digital dental radiography, comprising the following steps:
a) detecting first reference signals generated by the elements of a detector device, consisting of a CCD
array provided with a scintillator layer, when the detector device is not exposed to any x-radiation generated by a source of X-radiation;
b) detecting second reference signals generated by the elements of the detector device when the detector device is exposed to a known X-radiation from the source of X-radiation;
c) detecting third electrical signals generated by the elements of the detector device when the detector device is exposed to X-radiation representing an image of an object;
d) controlling the temperature of the detector device during steps a) to c) so as to maintain a constant temperature;
e) correcting the third electrical signals, detected in step c), on the basis of the first and second reference signals generated in steps a) and b) so as to compensate fluctuations in the third electrical signals of the individual elements of the CCD array caused by the dark current of these array elements, by the variation in the conversion efficiency of these array elements and by the inhomogeneity of the scintillator layer.
a) detecting first reference signals generated by the elements of a detector device, consisting of a CCD
array provided with a scintillator layer, when the detector device is not exposed to any x-radiation generated by a source of X-radiation;
b) detecting second reference signals generated by the elements of the detector device when the detector device is exposed to a known X-radiation from the source of X-radiation;
c) detecting third electrical signals generated by the elements of the detector device when the detector device is exposed to X-radiation representing an image of an object;
d) controlling the temperature of the detector device during steps a) to c) so as to maintain a constant temperature;
e) correcting the third electrical signals, detected in step c), on the basis of the first and second reference signals generated in steps a) and b) so as to compensate fluctuations in the third electrical signals of the individual elements of the CCD array caused by the dark current of these array elements, by the variation in the conversion efficiency of these array elements and by the inhomogeneity of the scintillator layer.
2. A method according to claim 1, wherein the first reference signals are detected in step a) by averaging several pictures of the detector device while this is not exposed to any X-radiation generated by the source of X-radiation.
3. A method according to claim 1 or 2, wherein the second reference signals are detected in step b) by averaging several pictures of the detector device while this is exposed to a known X-radiation from the source of X-radiation.
4. A method according to one of the claims 1 to 3, wherein the first reference signal of each element of the detector device is used as the offset value (O ij) for the element concerned and wherein a gain normalization value (G ij') is calculated for each element of the detector device on the basis of a nominal value of the electrical signals of the elements of the detector device at the known X-radiation and the first and the second reference signals.
5. A method according to claim 4, wherein the third electrical signals representing the image of an object are corrected by subtracting the offset value (O ij) for each element from the third electrical signal for the element concerned and by multipying the result of the subtraction by the gain normalization value (G ij') for each element.
6. A method according to claim 4 or 5, wherein the gain normalization value (G ij') for each element is calculated by subtracting the respective first reference signal from the respective second reference signal and by dividing the nominal value at the known X-radiation by the result of the subtraction.
7. A method according to one of the claims 1 to 6, wherein the known X-radiation is the maximum permissible X-radiation for the detector device.
8. A method according to one of the claims 1 to 7, wherein steps c) and e) are performed repeatedly and where the detector device is immersed in a fluid at a predetermined temperature between successive steps c).
Applications Claiming Priority (1)
| Application Number | Priority Date | Filing Date | Title |
|---|---|---|---|
| PCT/EP1997/005871 WO1999022252A1 (en) | 1996-04-17 | 1997-10-23 | Device and method for producing images in digital dental radiography |
Publications (1)
| Publication Number | Publication Date |
|---|---|
| CA2295893A1 true CA2295893A1 (en) | 1999-05-06 |
Family
ID=8166767
Family Applications (1)
| Application Number | Title | Priority Date | Filing Date |
|---|---|---|---|
| CA002295893A Abandoned CA2295893A1 (en) | 1997-10-23 | 1997-10-23 | Method for producing images in digital dental radiography |
Country Status (5)
| Country | Link |
|---|---|
| EP (1) | EP0974064B1 (en) |
| AT (1) | ATE198232T1 (en) |
| AU (1) | AU5313598A (en) |
| CA (1) | CA2295893A1 (en) |
| DE (1) | DE59702797D1 (en) |
Cited By (2)
| Publication number | Priority date | Publication date | Assignee | Title |
|---|---|---|---|---|
| WO2006108928A1 (en) * | 2005-04-12 | 2006-10-19 | Planmeca Oy | Ccd sensor and method for expanding dynamic range of ccd sensor |
| US8279315B2 (en) | 2005-04-12 | 2012-10-02 | Planmeca Oy | CCD sensor and method for expanding dynamic range of CCD sensor |
Families Citing this family (1)
| Publication number | Priority date | Publication date | Assignee | Title |
|---|---|---|---|---|
| DE102004016585B4 (en) | 2004-03-31 | 2006-02-09 | Siemens Ag | Method for noise correction in a flat panel detector |
-
1997
- 1997-10-23 AU AU53135/98A patent/AU5313598A/en not_active Abandoned
- 1997-10-23 AT AT97950023T patent/ATE198232T1/en active
- 1997-10-23 EP EP97950023A patent/EP0974064B1/en not_active Expired - Lifetime
- 1997-10-23 CA CA002295893A patent/CA2295893A1/en not_active Abandoned
- 1997-10-23 DE DE59702797T patent/DE59702797D1/en not_active Expired - Lifetime
Cited By (2)
| Publication number | Priority date | Publication date | Assignee | Title |
|---|---|---|---|---|
| WO2006108928A1 (en) * | 2005-04-12 | 2006-10-19 | Planmeca Oy | Ccd sensor and method for expanding dynamic range of ccd sensor |
| US8279315B2 (en) | 2005-04-12 | 2012-10-02 | Planmeca Oy | CCD sensor and method for expanding dynamic range of CCD sensor |
Also Published As
| Publication number | Publication date |
|---|---|
| EP0974064A1 (en) | 2000-01-26 |
| EP0974064B1 (en) | 2000-12-20 |
| AU5313598A (en) | 1999-05-17 |
| DE59702797D1 (en) | 2001-01-25 |
| ATE198232T1 (en) | 2001-01-15 |
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