|Publication number||US7504635 B2|
|Application number||US 10/548,985|
|Publication date||17 Mar 2009|
|Filing date||9 Mar 2004|
|Priority date||11 Mar 2003|
|Also published as||EP1601992A1, US20060180767, WO2004081606A1, WO2004081606A8|
|Publication number||10548985, 548985, PCT/2004/982, PCT/GB/2004/000982, PCT/GB/2004/00982, PCT/GB/4/000982, PCT/GB/4/00982, PCT/GB2004/000982, PCT/GB2004/00982, PCT/GB2004000982, PCT/GB200400982, PCT/GB4/000982, PCT/GB4/00982, PCT/GB4000982, PCT/GB400982, US 7504635 B2, US 7504635B2, US-B2-7504635, US7504635 B2, US7504635B2|
|Original Assignee||Symetrica Limited|
|Export Citation||BiBTeX, EndNote, RefMan|
|Patent Citations (20), Non-Patent Citations (6), Referenced by (1), Classifications (8), Legal Events (2)|
|External Links: USPTO, USPTO Assignment, Espacenet|
This application is a national phase of International Application No. PCT/GB2004/000982 filed Mar. 9, 2004 and published in the English language.
The invention relates to gamma-ray camera systems, in particular to spectral processing of data in gamma-ray camera systems.
The gamma-ray imager 14 includes a parallel collimator 22, a scintillator crystal 24, a light guide element 26 and a plurality of photo-multiplier tubes 28. The scintillator crystal is, for example, a large single crystal of Thallium doped Sodium Iodide (NaI(Tl)). The scintillator crystal is shielded from gamma-ray photons not incident through the parallel collimator by a shield 30.
The gamma-ray imager provides a 50 cm square image plane. The parallel collimator comprises an array of apertures with a characteristic cell size of 2 mm. Each cell provides a field-of-view of around 3°. The photo-multiplier tubes form a close packed hexagonal array of 61 tubes arranged to collectively view much of the 50 cm square scintillator crystal 24 forming the image plane.
During an exposure period, which in nuclear medicine imaging applications might typically be around 5 minutes, photons are emitted by the gamma-ray source in all directions. In a typical application, the photons will be emitted by radio-labelled pharmaceuticals in a patient's body (i.e. pharmaceuticals labelled with a radioactive tracer). In the example shown in
The signals from the photo-multiplier tubes are supplied to the detector read-out component. The detector read-out component determines the X- and Y-coordinates of the scintillation event from the relative intensities of the signals seen by each of the photo-multiplier tubes. The detector read-out component also calculates the total energy deposited in the scintillation event from a summation of the signal amplitudes seen by the photo-multiplier tubes. Read-out components for Anger-type gamma-ray cameras are well known . One mode of operation is known as list-mode operation. In this mode the detector read-out component outputs a signal in response to each scintillation event, the output signal including the calculated energy deposited in the scintillator crystal and the X- and Y-coordinates of the scintillation event. The output signals from the detector read-out component are coupled to the energy spectra accumulating component 12.
The functionality of the energy spectra accumulating component 12 in this example is provided by a suitably programmed general purpose computer. The computer includes an appropriate interface to receive and decode the output signals from the detector read-out component 16. The energy spectra accumulator 18 within the energy spectra accumulating component 12 operates to generate a three-dimensional observed data array I(X, Y, E). This array comprises a count of the number of scintillation events occurring within an exposure period as a function of their X, Y position within the detector, and the energy deposited in the scintillator crystal. A bin-size used in generating I(X, Y, E) is typically 5 mm for each spatial coordinates (i.e. X and Y) and 2 keV for the energy coordinate (i.e. E). Although the scintillator crystal 24 is a large single crystal, the 5 mm spatial binning used in generating I(X, Y, E) defines effective detector pixels, and these have a size of around 5 mm square. However, because of Compton scattering within the sample 4, the resolution in a resulting image is worse than this.
When an exposure is completed, the observed data array I(X, Y, E) is normalized to the exposure time and stored in the data storage component 20 of the energy spectra accumulating component for later analysis. I(X, Y, E) will typically be used in generating two-dimensional diagnostic images representing gamma-ray emission intensity from the source 4 as seen within selected gamma-ray energy ranges.
When forming a diagnostic image from the stored data array I(X, Y, E), for example in order to represent the distribution of Cobalt-57 within the field-of-view of the gamma-ray camera, the relatively poor intrinsic resolution is not necessarily a significant problem with an isolated point-like gamma-ray source. This is because for each pixel in the image, the intensity of gamma-ray emission seen by the gamma-ray camera can be represented by a summation over the full width of the peak around 122 keV, for example between 100 keV and 150 keV. Because the poor energy resolution does not directly effect the imaging capabilities of the camera in such cases, a summation over this wide energy range provides the best signal-to-noise ratio possible by making use of all detected events, without unduly compromising image quality.
In practice, however, it will generally be necessary to generate diagnostic images of gamma-ray emission within the field-of-view of the gamma-ray camera by summing the energy spectra recorded in each pixel over a more restricted energy range than the full width of the peak seen in
In cases where the field-of-view includes only a monochromatic source of gamma-ray emission, the poor energy resolution can still lead to degradation in derived diagnostic images where the source is embedded in a body. This is due to scattering, for instance Compton scattering, in the body surrounding the gamma-ray source.
The energy of gamma-rays emitted by radioactive sources used to label commonly-used pharmaceuticals in nuclear medicine is typically on the order of 102 keV. For example, Technetium-99 emitting at 140 keV is commonly used. This energy is chosen to be sufficiently energetic to allow emitted gamma-ray photons to escape from the surrounding body in which the gamma-ray source is embedded, yet without being so energetic as to make collimation and detection difficult. One disadvantage of this choice of energy is that the gamma-ray photons have a relatively high probability of scattering within the surrounding body before being viewed by the gamma-ray camera. An example of such a scattering event is shown by the photon labelled F in
In order to minimise the deleterious effect of Compton scattering in the source, it is necessary to reject as many of the Compton scattered detection events as possible, without unnecessarily discarding too many of the direct detection events. As noted above, when a gamma-ray photon is Compton scattered, some of its energy is imparted to a free electron in the scattering material. This means that, in principle, it is possible to distinguish between direct detection events and Compton scattered detection events on the basis of the energy deposited in the scintillator material comprising the scintillator detector. However, the poor energy resolution of scintillator crystal based gamma-ray camera systems is not able to provide suitable energy discrimination at 140 keV to make this an efficient process.
On the high energy side of the peak seen in
As a consequence of the failings of gamma-ray camera systems of the type described above, alternative designs for gamma-ray camera systems which help to improve the rejection of Compton scattered detection events while maintaining the direct detection events have evolved.
For example, a number of research groups are seeking to exploit the improved energy resolution characteristics of room-temperature semiconductor detectors such as CdZnTe. These are capable of providing an intrinsic full width at half maximum energy resolution of around 4% at 140 keV . This compares favourably with the approximately 20% seen above for an Anger-type gamma-ray camera system. The improved energy resolution allows a narrower energy window to be employed whilst retaining many more of the direct detection events since these are now contained in a narrower spectral peak. However, there are a number of disadvantages associated with this type of design. For example, each detector pixel requires a separate read-out electronics channel, the detection efficiency is relatively low and the camera costs are high.
Improved energy resolution has also been achieved in the recently introduced “2020tc Imager” camera manufactured by Digirad, San Diego, USA. This uses silicon positive-intrinsic-negative (PIN) diodes to record the scintillation flash generated in a large array of discrete Caesium Iodide (CsI) scintillation crystal elements. However, this design again requires significantly increased complexity in the read out electronics in comparison to other types of gamma-ray camera system, such as the kind discussed above.
Methods of improving the spectral resolution of non-imaging scintillator-based gamma-ray detectors by post-exposure spectral processing have also been described [3, 4, 5]. These techniques involve determining a detector response function for the single element detector and deconvolving this response function from observed spectra. However, these techniques have been applied only to single element specific scintillator geometries so as to provide a calculable transfer function describing the response of a photo-detector to pulses of optical photons from a scintillation event.
Various other schemes for reducing the effect of Compton scattered detection events in gamma-ray camera systems have been attempted.
U.S. Pat. No. 5,903,008 describes the use of dual energy-windows for diagnostic image formation . This dual energy-window technique measures the relative number of counts recorded in two energy channels; one centred on the photo-peak energy and the second, some 10-15% below that energy. This ratio is measured initially when there is no scattering present. Thereafter, the photo-peak values are modified according to the value of the number of counts in the lower window. This value is dominated by the presence of scattered events.
U.S. Pat. No. 5,530,248 describes a scheme for theoretical modelling of the contribution of Compton scattered detection events to the energy spectra . This technique generates a trial function from the energy spectrum recorded in each pixel when no scattering material is present. This is achieved by taking the first differential of the spectrum to emphasise the photo-peak. This modified function is then fitted to the energy spectra seen when a scattering medium is present in order to distinguish between ‘wanted’ and ‘unwanted’ photons.
U.S. Pat. No. 5,633,499 describes a scheme based on calculating a correction table to be applied to an image acquired using a conventional method of selecting only those events that fall within an energy window spanning ±10% of the photo-peak . The correction value is derived from the measurement of the centroid of the spectrum recorded in a particular detector pixel. The data are used to estimate the scatter contribution.
U.S. Pat. No. 5,561,297 describes a spectrum subtraction method . This method in essence uses the idea of subtracting a reference spectrum for each detector pixel form an observed spectrum. The subtracted reference spectrum is that recorded in the absence of a scatterer.
According to a first aspect of the invention there is provided a gamma-ray camera system comprising: a gamma-ray imager including detector pixels formed from a scintillator material; a detector read-out component for determining positions and energies of scintillation events occurring within the scintillator material; an energy spectra accumulating component for compiling observed energy spectra of scintillation events in the detector pixels; and a spectra processing component operable to deconvolve a pixel specific detector response function from the observed energy spectra.
By deconvolving detector response functions which are dependent on individual detector pixels from the observed energy spectra, much higher energy resolution can be obtained when compared to previously described systems. The higher energy resolution allows, for example, improved images to be formed where there is significant Compton scattering in a source being imaged. The higher energy resolution also allows radioactive sources emitting gamma-ray photons at different, though similar, energies to be better distinguished. Whereas with previous systems, it has generally been necessary to make sequential observations of a subject in order to study how different pharmaceuticals marked with different radioactive tracers are taken up in a body, improved energy resolution allows different radioactive tracers to be simultaneously imaged and separated in a single observation.
In one example, the detector response functions are selected from a store of detector response functions which include detector response functions specific to individual detector pixels.
By providing a separate detector response function for each individual detector pixel, the most accurate mapping of the detector response to detector pixel can be obtained.
Alternatively, the detector response functions may be selected from a store of detector response functions which comprises a family, and for the pixels in the gamma-ray image a detector response function is selected from the family.
By providing a number of detector response functions which is fewer than the number of detector pixels, a compromise can be made between the benefits of providing individual detector response functions for each detector pixel, and the benefits of working with a reduced number of detector response functions, for instance reduced storage requirements.
The invention is equally applicable to a range of gamma-ray camera systems employing different detection planes. For instance, an Anger-type gamma-ray imager in which the scintillator material comprises a single scintillator crystal optically coupled to a plurality of photo-detectors may be used. Alternatively, other types of gamma-ray imager may be employed, for example, gamma-ray imagers in which the scintillator material comprises an array of discrete scintillator crystal elements optically coupled to a corresponding array of discrete photo-detectors.
According to a second aspect of the invention there is provided a method of processing energy spectra recorded in detector pixels of a gamma-ray scintillation camera comprising: providing observed energy spectra recorded in the detector pixels; and deconvolving a pixel specific detector response function from the observed energy spectra.
This method provides much higher energy resolution than previous methods of processing data from gamma-ray camera systems and correspondingly provides similar benefits to those of the first aspect of the invention described above.
The method of processing energy spectra may include selecting detector response functions from a store of detector response functions which include a detector response function specific to individual detector pixels.
Alternatively, the method may include selecting detector response functions from a store of detector response functions which comprises a family, and for the detector pixels a detector response function is selected from the family.
A computer program product bearing machine readable instructions for implementing the method is also provided.
According to a third aspect of the invention there is provided a method of generating pixel-specific detector response functions for a gamma-ray imager including detector pixels formed from a scintillator material comprising: obtaining a theoretical component to the detector response as a function of position by performing a simulation; obtaining an empirical component to the detector response as a function of position by observing the response of the gamma-ray imager to a calibration source; and combining the theoretical and empirical components to provide detector response functions for the detector pixels.
A computer program product bearing machine readable instructions for implementing the method and a data storage medium bearing detector response functions calculated according to the method are also provided.
For a better understanding of the invention and to show how the same may be carried into effect reference is now made by way of example to the accompanying drawings in which:
The spectra processing component 34 is arranged to read the observed data array I(X, Y, E) from the data storage component 20 in the energy spectra accumulating component 12. When an exposure is completed and an observed data array I(X, Y, E) generated as described above, the observed data array is read by the spectra processor 38 in the spectra processing component 34. The spectra processor is operable to deconvolve a detector response function from the individual spectra associated with each pixel in the observed data array I(X, Y, E), the specific detector response function employed in the deconvolution is selected from a store of detector response functions 36. The choice of detector response function is based on which detector pixel the spectrum currently being processed is associated with. Once each of the energy spectra comprising the observed data array I(X, Y, E) have been deconvolved, the resulting refined energy spectra are collated into a refined data array S(X, Y, E). The refined data array is of the same form as I(X, Y, E) and may be used in the same way to generate diagnostic images.
With a gamma-ray camera system of the type shown in
where the detector response function is parametrized such that RX
The much improved resolution makes it easier to distinguish Compton scattered detection events from direct detection events when forming diagnostic images from the refined data array S(X, Y, E). This is possible since a narrower energy widow around the peak energy may be used when generating diagnostic images so as to discard much of the Compton scattered detection events, while maintaining most of the direct detection events. For instance, the width of the narrow energy window W2 shown in
The difficulties in applying this technique to gamma-ray camera systems are the complications introduced by the fact that the response of a relatively large single scintillator crystal is not only a function of photon energy, but it is also a function of the location of the gamma-ray interaction. This means that it is necessary to construct a model of the detector response function not only for the way that gamma-rays interact in the scintillator crystal material itself, but it is also necessary to modify this model with reference to an empirical calibration of the overall response of the gamma-ray camera system on a detector pixel by detector pixel basis. The detector response function varies from one detector pixel to another as a consequence of non-uniformities in the scintillator crystal light-yield, and also in the light-collection efficiency on the gamma-ray camera imager. For example, the light collection efficiency will depend on detector pixel both due to differences in transfer function from different scintillation sites to the photo-multiplier tubes, and also to non-uniformities in the response of the photo-multiplier tubes. Further variations in the detector response function for different detector pixels are introduced as a consequence of non-linearities in the response of the scintillation crystal as a function of energy deposited in a scintillation event.
By using, for example, a Monte Carlo technique, it is possible to predict the way that the scintillator material responds to incident gamma-ray photons at energies within an energy range of interest, for instance between 50 keV and 500 keV. Other modelling methods could also be used. This information may then be combined with empirical calibration data to provide information on how the gamma-ray imager actually responds at a number of discrete energies within the same region of interest. The calibration data may be acquired, for example, by observing a number of monochromatic radioactive sources emitting within the energy range of interest. By experimentally determining the response of each individual detector pixel to known incident gamma-ray energy spectra, the positional dependence of the detector is response function can be determined. With a point calibration source, data can be obtained for each pixel simply by removing the collimator and placing the source at a distance of say 50 cm from the scintillator crystal, or by mounting the calibration source on a translation stage such that it can be scanned across the field-of-view of the camera. It is most appropriate to determine how the detector response function varies with position at a spatial resolution comparable with that of the gamma-ray imager. The determined detector response functions are specific to an individual gamma-ray camera.
It will also be appreciated that while in the above examples a Monte Carlo or similar simulation provides a base model which is modified according to empirically determined calibration data, in other examples the simulation may be dispensed with. In such cases purely empirical pixel specific detector response functions can be obtained from observations of the response of the gamma-ray imager to point calibration source.
With accurate models for the way in which the gamma-ray imager responds as a function of incident gamma-ray photon energy, deconvolution can be performed on a detector pixel by detector pixel basis using the techniques noted above . Accordingly, following an exposure, when energy spectra for each detector pixel have been recorded in an observed data array I(X, Y, E) as described above, the gamma-ray spectrum incident on each detector pixel can be recovered using the position-sensitive detector response functions in a standard deconvolution algorithm.
Because there are more than 2000 detector pixels present in the gamma-ray camera system, the process shown in
In other examples, a combination of the methods shown in
Whilst in the above examples, individual detector response functions are determined and stored for each detector pixel, it will be appreciated that in some circumstances the variation of detector response on spatial scales comparable to the detector pixel size will be small. In such cases, it may be unnecessary to store a detector response function for each pixel. For instance, in a gamma-ray camera system with a 50 cm square scintillator crystal detector and an imaging resolution of 5 mm, detector response functions may be determined with a spatial resolution of, for example, 1 cm. In this way a family of detector response functions are calculated corresponding to an array of 1 cm square elements spanning the scintillator crystal. When selecting a suitable detector response function to deconvolve form the observed data array corresponding to detector pixel at X, Y, the detector response function corresponding to the 1 cm square element which includes detector pixel X, Y is chosen.
While the above described examples of gamma-ray camera systems have used Anger-type gamma-ray cameras, it will be appreciated that in other embodiments spectra processing components of the type described above may be used in conjunction with other gamma-ray camera types.
As noted above, the detector response functions associated with the different detector pixels in a gamma-ray camera system are specific to each particular gamma-ray camera system. Accordingly, appropriately calculated detector response functions may be provided along with a gamma-ray camera system at first supply to an end user. In cases where the functionality of the spectra processing component is to be provided by a general purpose computer, the calculated detector response functions and software operable to configure the general purpose computer may be supplied together or separately on a computer program product, for instance a CD-ROM or other data-storage medium. In cases where the functionality of the spectra processing component is to be provided in firmware or hardware, for example by an ASIC, a FPGA or a DSP, the calculated detector response functions can, for example, be stored in ROM on an integrated chip along with the firmware.
In addition to being shipped together at first supply, spectra processing components similar to those described above can be coupled to existing gamma-ray camera systems as an ‘after-market’ add-on for improving energy resolution. To provide the best possible improvements in energy resolution, an existing user of a gamma-ray camera system may provide a third party with access to the camera system, such that the third party can fully calculate the detector response functions using the techniques described above. Alternatively, the existing user might supply the third party with appropriate calibration data from the gamma-ray camera from which the third party calculates the appropriate detector response functions to be supplied to the existing user. As above, the detector response functions may be supplied alone, for example on a data-storage medium, or along with a computer program product bearing machine readable instructions for implementing the functionality of a spectra processing component. In other cases, a hardware add-on may be supplied to an existing user. The hardware add-on including, for example, an appropriately configured interface for interfacing with the detector read-out component of the existing gamma-ray camera system, firmware or hardware, for example an ASIC, a FPGA or a DSP, for providing the functionality of a spectra processing component, and a memory, for example ROM or a replaceable data-storage medium, for storing appropriate detector response functions.
While in the above examples, the detector response functions are calculated for individual gamma-ray camera systems, improvements in spectral resolution may also be achieved by deconvolving more generic pixel dependent detector response functions. For example, for commonly used configurations of gamma-ray camera system, detector response functions may be calculated for each detector pixel in a manner dependent only on properties of the scintillator material and the geometric configuration of the gamma-ray camera imager.
This could be done, for example, by performing a Monte Carlo modelling simulation of a commonly used gamma-ray camera imager configuration without subsequently modifying the model according to specific calibration data obtained with an individual example gamma-ray camera imager. By disregarding effects which are specific to individual gamma-ray camera imagers, for example non-uniformities in the response of the actual scintillator material and/or photo-detectors employed, more widely applicable detector response functions may be provided. While the use of more generic detector response functions does not provide the maximum possible improvement in energy resolution, it does provide a way to quickly and easily improve the performance of a large number of similarly configured gamma-ray systems without requiring individual calibration.
Another way to obtain more generic detector response functions for a particular gamma-ray camera imager configuration would be to determine the average of a number of previously determined detector response functions seen in examples of the imager configuration of interest.
|Cited Patent||Filing date||Publication date||Applicant||Title|
|US3944835 *||25 Sep 1974||16 Mar 1976||General Electric Company||High energy radiation detector having improved reflective backing for phosphor layer|
|US4839808||22 May 1987||13 Jun 1989||The University Of Michigan||Correction for compton scattering by analysis of energy spectra|
|US5530248||30 May 1995||25 Jun 1996||Elscint Ltd.||Method of and apparatus for improving gamma camera image|
|US5561297||12 Jul 1993||1 Oct 1996||Sopha Medical Systems, Inc.||Scatter correcting gamma-ray camera|
|US5576546 *||19 Apr 1994||19 Nov 1996||Park Medical Systems Inc.||Depth-of-interaction normalization of signals for improved positioning, and energy resolution in scintillation camera|
|US5633499||21 Nov 1995||27 May 1997||Trionix Research Laboratory, Inc.||Scatter elimination technique and apparatus in radionuclide emission and transmission imaging in a nuclear camera|
|US5692507 *||7 Dec 1994||2 Dec 1997||Varian Associates, Inc.||Computer tomography apparatus using image intensifier detector|
|US5903008||2 Jul 1997||11 May 1999||General Electric Company||Scatter correction methods and systems in single photon emission computed tomography|
|US6057551 *||8 Sep 1997||2 May 2000||Smv International||Process for processing pulses delivered by a gamma camera and a gamma camera using this process|
|US6194715 *||5 Apr 1999||27 Feb 2001||Digirad Corporation||Semiconductor gamma-ray camera and medical imaging system|
|US6232954 *||8 May 1997||15 May 2001||Imation Corp.||Arrangement for high-accuracy colorimetric characterization of display devices and method therefor|
|US6369389||24 Nov 1996||9 Apr 2002||Ge Medical Systems Israel, Ltd.||Real-time compton scatter correction|
|US6628744 *||26 Sep 2002||30 Sep 2003||Koninklijke Philips Electronics N.V.||Off-focal radiation correction in CT|
|US20010001011 *||29 Dec 2000||10 May 2001||Jesse Salb||System and method for radiographic imaging of tissue|
|US20020011571 *||30 Apr 2001||31 Jan 2002||Lin Gregory Sharat||Compton deconvolution camera|
|US20020036269 *||30 Jul 2001||28 Mar 2002||Imarad Imaging Systems Ltd.||Gamma-ray detector for coincidence detection|
|US20020051111 *||17 Dec 2001||2 May 2002||Greene Raymond G.||Construction of large, robust, monolithic and monolithic-like, AMLCD displays with wide view angle|
|US20050055174 *||4 Dec 2000||10 Mar 2005||V Target Ltd.||Radioactive emission detector equipped with a position tracking system and utilization thereof with medical systems and in medical procedures|
|EP1004983A1||5 Nov 1999||31 May 2000||General Electric Company||High resolution radiation imaging system|
|WO1998023973A1||24 Nov 1996||4 Jun 1998||Elscint Ltd||Real-time compton scatter correction|
|1||Anger, Hal O. "Scintillation Camera." The Review of Scientific Instruments. vol. 29, No. 1, Jan. 1958.|
|2||Berger, M.J. and S.M. Seltzer. "Response Functions for Sodium Iodide Scintillation Detectors." Nuclear Instruments and Methods. Vol. 104 (1972), pp. 317-332.|
|3||Meng, L.J. and D. Ramsden, "An Inter-comparison of Three Spectral-Deconvolution Algorithms for Gamma-ray Spectroscopy", IEEE Transaction on Nuclear Science, vol. 47, No. 4, Aug. 2000, pp. 1329-1336.|
|4||Meng, L.J. et al. "'Scintisphere'-The Shape of Things to Come in Gamma-Ray Spectroscopy." IEEE Transactions on Nuclear Science. Vol. 49, No. 4, Aug. 2002, pp. 1681-1686.|
|5||Meng, L.J. et al. "The design and performance of a large-volume spherical Csl (T1) scintillation counter for gamma-ray spectroscopy." Nuclear Instruments and Methods in Physics Research. A 485 (2002), pp. 468-476.|
|6||Parnham, K.B., et al. "Development, Design and Performance of a CdZnTe-Based Nuclear Medical Imager." Proceedings of the SNM 49th Annual Meeting. vol. 43, No. 5, May 2002.|
|Citing Patent||Filing date||Publication date||Applicant||Title|
|US20140084149 *||27 Sep 2012||27 Mar 2014||Christian Stoller||Use of spectral information to extend temperature range of gamma-ray detector|
|International Classification||G01T1/20, G01T1/164|
|Cooperative Classification||G01T1/1644, G01T1/1642, A61B6/4258|
|European Classification||G01T1/164B1, G01T1/164B2|
|19 Oct 2005||AS||Assignment|
Owner name: SYMETRICA LIMITED, UNITED KINGDOM
Free format text: ASSIGNMENT OF ASSIGNORS INTEREST;ASSIGNOR:RAMSDEN, DAVID;REEL/FRAME:016657/0846
Effective date: 20050817
|22 Aug 2012||FPAY||Fee payment|
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