a system of the present invention possess a wide spectrum of properties to be used effectively both in implantology and for the applications described . the properties are summarized below in table 1 . none of the previous dentomaxillofacial imaging modalities possesses all these properties , and the practitioner has historically been face with various trade - offs in the decision as to which modality to use . the new system features all the properties from table 1 , and is therefore expected to become the modality of choice in dental implantology and other clinical tasks that involve the dentomaxillofacial complex . a preferred system of the invention is generally indicated at 10 in fig1 a . its building blocks are : 1 ) a cone - beam data acquisition and image reconstruction devices including a computer 12 with a display 14 ; 2 ) a detector assembly , generally indicated at 16 , including an amorphous silicon ( asi : h ) detector array coupled to a source assembly , generally indicated at 18 , including an x - ray source , both assemblies being mounted on a rotating gantry , generally indicated at 20 , and fitted with appropriate x - ray filtration , collimation , shadow masks , etc . ; 3 ) the computer 12 or data processor is programmed with a ) a penalized weighted least squares ( pwls ) reconstruction program and / or b ) a dual energy penalized weighted least squares ( de pwls ) reconstruction program when precise information regarding soft tissue and bone densities is needed . compared to the conventional fan - beam geometry , a cone - beam geometry is used in the present invention due to its high efficiency in x - ray use , inherent quickness in volumetric data acquisition , and potential for reducing the cost of ct machines . conventional fan - beam scans are obtained by illuminating an object with a narrow , fan - shaped beam of x - rays . the x - ray beam generated by the tube is focused to a fan - shaped beam by rejecting the photons outside the fan , resulting in a highly inefficient use of the x - ray photons . further , the fan - beam approach requires reconstructing the object slice - by - slice and then stacking the slices to obtain a 3 - d representation of the object . each individual slice , therefore , requires a separate scan and separate 2 - d reconstruction . the cone - beam technique , on the other hand , requires only a single scan to capture the entire object with a cone of x - rays 22 . the time required to acquire a single cone - beam projection is the same as that required by a single fan - beam projection . but since it takes several fan - beam scans to complete the imaging of a single object , the acquisition time for the fan - beam tends to be much longer than with the cone - beam . this is important because the more time the patient is subjected to the scanning process , the more likely it is that the patient may move or swallow , blurring the scan . although it may be possible to reduce the acquisition time of the fan - beam method by using a higher power x - ray tube , this increases the cost and bulkiness of the scanner . the imaging detector 16 used in the system 10 is a large - area array constructed from hydrogenated amorphous silicon ( asi : h ) as described in the above - noted application . the detector 16 is a “ self - scanned ” array of n - i - p photodiodes and thin - film transistor switches . the integrated - flux mode x - ray detector 16 is constructed by mating a flat - panel photodiode array with the appropriate scintillator . the detector 16 is both inexpensive and capable of generating high - quality images . essentially the same technology is used to construct active - matrix , flat - panel computer displays and large - area document imagers . they are replacing film and image - intensifiers in conventional radiography and fluoroscopy applications . several companies currently produce detector arrays for commercial sale . several years ago it was predicted that these devices could be used for tomography and for attenuation correction applications in pet and spect . although it has been recently suggested that a special image intensifier can be coupled with a ccd camera for use in a cone - beam - based dental ct scanner , the characteristics of asi : h - based flat panel detector arrays make them a much better choice for this application . the asi : h detector arrays offer three distinct advantages . first , the asi : h flat panel detectors , unlike image intensifiers , do not create geometric distortions that must be addressed when processing the data . second , flat panel detectors are available in sizes up to 40 × 40 cm , which is large enough to cover the entire head , whereas image intensifiers have relatively smaller diameters , creating “ truncated view ” artifacts . third , flat panel detectors afford a greater dynamic range than that offered by the image intensifier + ccd camera approach . the a - si photodiode array / scintillator combination is not the only type of flat - panel imaging detector that can be used in the system 10 described here . in particular , direct detection arrays , which have an x - ray converter such as lead iodide or amorphous selenium ( or other suitable converter ), can be used . these devices have no intermediate light conversion step and instead convert x - rays to an electrical charge that is read out by the array . although direct detection devices may eventually offer advantages such as higher spatial resolution , they are at present investigational having detection efficiency that is too low at the energies used in tomography applications . the vertically - oriented , high - resolution tomograph system 10 is primarily provided for dentomaxillofacial imaging applications . the basic system 10 includes a support structure , generally indicated at 11 , including the gantry 20 having an arm 21 , which , in this case , adjusts to accommodate different size patients and a rotating arm ( as shown in fig1 c , 2 and 3 at 24 ) to which is affixed an x - ray source 26 and a controlled source collimator 28 ( i . e ., fig2 ) of the assembly 18 on one end of the arm 24 and on the other end of the arm 24 , the assembly 16 including a 2 - d position - sensitive x - ray detector 30 and a scatter - reducing collimator 32 . the system 10 also includes a device , generally indicated at 34 , for moving the detector 30 with respect to the arm 21 to accommodate different patient head sizes and to reduce the effect of detector non - uniformities . the device 34 includes a detector translation track 35 . in order to acquire images that quantitatively assess the quality ( i . e ., mineral content ) of bone , the instrument can switch between two x - ray tube potentials or an x - ray filter 36 can be interposed that spatially modulates the source ( i . e ., a “ shadow mask ”). at each rotation angle of the gantry arm 21 , the x - ray source 26 irradiates the object with two different x - ray spectra and two images are recorded from the detector 30 . more detailed drawings of the assemblies 16 and 18 are shown in fig2 and 3 wherein the detector 30 is translated along the track 35 to various offset positions and the position of the detector 30 controls the position of the source collimator 28 by the use of the mechanical arm 24 . the purpose of these features are discussed hereinbelow . to accomplish scanning , the patient sits upright in a chair 40 of the support structure 11 . the chair may be vertically adjustable . the patient may , for purposes of reducing head - motion during the scan , bite into an immobilizing apparatus 42 ( i . e ., fig1 c ) affixed to the tomograph ( e . g ., a dental impression tray filled with impression material as used in conventional linear tomography ). during set - up for the scan , the operator preselects the desired x - ray tube ( i . e ., source ) potential ( kv ) and the instrument performs a “ scout scan ” as described in more detail below . from the scout scan , not only can the optimum exposure time for each frame be estimated ( as it is done in current practice ), but the appropriate detector offset can also be computed . in order to acquire a dataset that is tomographically complete , the offset allows use of a detector that is smaller than would be ordinarily necessary to collect a tomographically complete dataset of the entire head ( complete with respect to the “ fan - beam ” slice ). it also provides the desirable added feature of reducing detected scatter as described below . although the instrument may utilize a device to stabilize the motion of the patient &# 39 ; s head , such as the bite plate , as illustrated in fig1 c at 42 , it is nonetheless expected that the patient will still slightly move during the scan ( which will typically last between 30 and 90 seconds ). this residual head motion will cause the data acquired at different angles to be inconsistent with each other , thus introducing artifacts in the reconstructed images . these artifacts , referred to as “ patient motion artifacts ,” typically result in images that are blurred . in its mild form , the blurring effectively reduces the spatial resolution of the device , while in severe cases it can render the image useless , and needs to be corrected for as the spatial resolution is at premium in the imaging of the dentomaxillofacial complex . a solution for correcting the head motion artifacts will now be described . the method consists of two parts : 1 ) a device that measures the relative motion of the head at each projection ; and 2 ) software method that incorporates these measurements into image reconstruction . the head is treated as a rigid body with six degrees of freedom . the motion of the head is measured with a device depicted in fig1 c . it consists of a sensor 44 that records the motion of the head . the recordings of the device are synchronized with the image acquisition sequence of the detector 30 so that for each detector position , the position of the head is known . these measurements are then fed into the computer 12 programmed to perform the image reconstruction method . the movement of the head in the coordinate system of the scanner can be looked at as the ( equivalent in magnitude but opposite in direction ) movement of the scanner ( source 26 and detector 30 ) in the coordinate system of the head . the expressions for all image reconstruction algorithms ( e . g ., filtered back projection , statistical image reconstruction methods , etc .) are typically written in the coordinate system of the reconstructed object ( in our case the head ) and involve the coordinates of the x - ray source 26 and detector 30 in this coordinate system . in the proposed method , these coordinates of the x - ray source 26 and detector 30 are different at each angle , where the difference corresponds to the measured head motion . in addition to single - energy scanning , where as in conventional instruments attenuation measurements through the object are taken using a single x - ray spectrum , the system of the present invention is capable of acquiring and processing data using two x - ray spectra . the advantages are two . first , it is well known that the technique can provide superior corrections for “ beam - hardening ” or the fact that the effective energy of the broadband x - ray bremsstrahlung radiation increases as the x - ray beam traverses soft tissue and bone and the lower energy x - rays are preferentially absorbed . second , the method , when combined with the appropriate image reconstruction technique ( e . g ., penalized weighted least squares ), can provide good estimates of bone - mineral contact as a measure of bone quality ( which is especially important in dental implantology ). there are numerous methods for accomplishing imaging using two spectra . a potentially important technique can generate two spectra by spatially modulating the x - ray beam using a “ shadow mask .” this device , a filter 36 that spatially modulates the x - ray beam , has been used in order to acquire dual - energy data simultaneously . the shadow mask 36 used here , however , can contain both filtration elements for dual - energy data acquisition as well as “ beam - stops ” for estimating residual scatter ( see below ). to acquire a set of cone - beam projection data that satisfies the cone - beam completeness conditions , the focal - spot of the x - ray source should intersect every plane through the object . obviously , a circular orbit does not satisfy these conditions ( although usable ct images have been obtained with circular orbits ). there are a number of simple orbits that can be used with the described device to satisfy these conditions . as two examples , the source 26 and the detector 30 can move axially during rotation to accomplish helical cone - beam scanning or the source 26 and the detector 30 can “ wobble ” to such an extent that the focal - spot of the source 26 intersects every plane passing through the desired axial extent of the scan ( fig3 ). this additional motion can be accomplished in a variety of ways from using an additional motor to move the gantry 20 vertically or wobble it during rotation by using a mechanical cam arrangement ( no shown ) that transfers some of the torque of the motor used to rotate the gantry 20 into the appropriate force . fig3 shows an optional pivot 50 on vertical axis of rotation 51 for gantry movement as indicated by arrows 52 . alternatively , the gantry 20 moves vertically as indicated by arrows 54 . the system 10 is capable of scanning a head along an elliptical orbit , as shown in fig1 b . an elliptical orbit is better suited for head imaging as the head is typically elliptical in shape . this is of particular interest when an offset detector arrangement is used as described herein . the elliptical arrangement allows the detector 30 to capture more data while moving along the sides of the head in comparison to a circular scan . the data acquired in this fashion is then fed into the data processor or computer 22 programmed with a modified image reconstruction program . using a “ scout - scan ” to estimate exposure and determine best detector offset it is common practice to use a low - current scan at the desired x - ray tube potential in order to determine whether patient positioning is proper and to estimate an appropriate exposure . in addition to using the scout scan for these purposes , the system 10 of the invention also uses the scan to determine an appropriate offset for the detector 30 in order to ( 1 ) ensure that a complete dataset is acquired , ( 2 ) ensure that enough information is available to calculate detector positions such that the tradeoff between x - ray scatter and improved information resulting from measuring some rays twice is appropriate . using an offset detector to reduce necessary size , improve sampling and decrease x - ray scatter the system 10 is capable of using a detector having a width smaller than that ordinarily required to obtain a complete tomographic dataset . typically , in fan - beam and cone - beam tomography , the x - ray image or projection of the object in each view must encompass the entire object . this , depending on the system geometry , can require use of a detector that is quite large , which increases system cost . it is well known that a detector of smaller size can be used . specifically , a detector half of the width can be used along with a single rotation plus an angle equivalent to the cone - angle to acquire a dataset that is essentially complete ( in the sense that it contains the same data contained in the scan using a full - size detector for the central slice ). data from the half - detector geometry can be reconstructed using a modified image reconstruction algorithm . in addition to cost - savings of using such an arrangement , the detector 30 can be combined with source collimation to irradiate only the portion of the object containing line segments connecting the source 26 with the detector 30 . this arrangement , as noted below , will reduce x - ray scatter . rather than using a detector having a fixed offset , as has been used previously , the detector 30 in this system 10 can move along the track 35 ( fig2 wherein a field of view of the system 10 is indicated at 46 ) to achieve a variable offset in order to collect a complete dataset . the offset of the detector 30 can either ( 1 ) be predetermined and set once per scan by the operator , or ( 2 ) be moved by motor control to positions determined from the scout scan or from previous frames of the diagnostic scan . the advantage of this approach is that the detector 30 can be optimally positioned for each object scanned . for example , the detector 30 may be of such a size that projections of the object “ almost ” fit the width . in this case , the detector 30 may not need to be displaced by a half - detector width . the advantage for doing so is that some portions of the object will be sampled twice by the x - ray source 26 and the detector 30 during the scan reducing noise due to the quantum nature of x - ray detection . fig4 - 6 show examples of the sinogram sampling pattern of the fan - beam slice for detector offsets ranging from none to one half - detector width . when the detector 30 has no offset , the entire sinogram is sampled twice . at full offset , it is only sampled once per rotation . at positions between these extremes , a portion of the object ( generally the thickest part ) is sampled twice , which can reduce noise . scattering — both compton and coherent — is the most prevalent interaction mechanism of x - ray photons with tissues of the body . while scattering contributes significantly to the observed attenuation of a beam of x - ray photons through the body , a large fraction of the scattered photons escape the body and are subsequently detected at the x - ray detector . recording these photons , whose direction has been altered considerably from the incident x - ray beam , leads to serious reductions in the contrast - resolution of x - ray ct . in conventional fan - beam or spiral - scan tomography , compton - scatter is reduced by two methods : 1 ) the x - ray source is finely collimated so that it only illuminates a narrow region subtended by the strip - detector ( either single - or multi - slice ) used in these systems , and 2 ) additional detector collimation may be used to reduce scatter . the additional collimation may either be a detector slit or a set of fan - beam channels , which in older xenon gas - based detectors were actually part of the detector . in cone - beam tomography , scatter can be much more severe because the source irradiates a much larger section of the object . therefore , scatter comes not only from a narrow axial region ( or slice ) but also from the entire irradiated axial extent . in conventional film radiography , scatter can be reduced by using a variety of scatter rejection grids . these grids are usually comprised of alternating strips of aluminum and lead foil . the strips are focused linearly to the focal spot of the x - ray source and reduce the detection of scattered radiation because , as with the use of collimation for spiral and conventional ct , the ray - path of scatter will typically not come a direction consistent with the x - ray source and will therefore be absorbed by one of the lead strips . the problems with using these grids for cone - beam tomography is that their transparency — the ratio of the exiting flux to the incident x - ray flux — is relatively low due to both a large number of lead strips ( typically 85 or 103 per inch ), which must be relatively thick to absorb x - rays of the appropriate energies , and the aluminum used as the interspacer . an effective scatter rejection grid with much higher transparency can be constructed with lead foil and machinable polyacrylimide foam — a technique that has been used for constructing slice collimators for spect . the basic construction is similar to the conventional grid for radiography described above : each grid or collimator 32 is a stack of absorbing septa 31 , focused on the focal - spot of the x - ray source assembly 18 , separated by a low - attenuation spacer 34 as shown in fig7 a and 7 b . in this case , however , the layers 31 of the scatter rejection grid 32 are much farther apart ( perhaps a maximum of 10 foil - foam layers 31 per inch rather than 85 ). the ratio of the grid , or the ratio of the depth of the linearly focused absorbing septa 31 to the space between septa , can be the same ; therefore , the scatter rejection properties can be similar but the transparency much higher ( the attenuation of foam is much smaller than aluminum ). two linearly focused grids placed orthogonally with respect to one another can be used to reduce scatter to an even greater degree . another method for constructing the crossed grids is to cut slots halfway through each absorbing septum in order to accept corresponding slots cut into orthogonal septa . this results in a largely self - supporting grid ( which can be placed within a rigid frame around the perimeter for stability ) requiring no foam supports . an additional advantage is that the depth of the grid is reduced by a factor of two . yet another technique involves using appropriately corrugated absorbing strips ( with corrugations focused on the focal spot of the x - ray source ) that are tapered in two dimensions such that the corrugations are focused on the x - ray source . these corrugated septa are then separated by planar septa . this technique is well known in the construction of collimators for nuclear medicine gamma cameras . other techniques used in nuclear medicine collimator construction such as casting can also be employed to obtain high - transparency , coarse scatter - rejection grids . the larger the extent of the irradiated portion of the object , the greater the scatter from the object . it is highly desirable from the viewpoints of both reducing patient dose and scattered x - rays that the source only irradiates the necessary portion of the object . in conventional radiography systems this is accomplished by using an adjustable source collimator to restrict the radiation field to the desired region of the object . as noted above , the system described here can use a detector having a width smaller than necessary for acquisition of a complete tomographic dataset . when used in conjunction with a source collimator that only irradiates the portion of the object “ seen ” by the detector less compton scatter will be generated ( and detected ) than in a system using a full - size detector . as shown in fig2 and 3 the detector 30 , as it moves along its track 35 , is connected to the x - ray source collimator 28 through the control arm 24 that ensures that the source 26 only irradiates portions of the object that are visible to the detector 30 . a significant fraction of photons will scatter multiple times within the object and many of these can pass the scatter - rejection grid 32 . to reduce the average effect of detected scatter on reconstructed images , it is desirable to employ a method that measures the distribution of residual scatter . once the mean distribution has been estimated , it can be removed from the data . the residual detected scatter can be measured by placing a series of beam - stops or shadow masks ( i . e ., shadow mask 36 ) between the x - ray source 26 and the object . in the absence of scatter , the shadow of the beam - stops on the detector 30 should record zero photons . photons that scatter in the object and that pass the scatter rejecting grid 32 will , however , result in some signal in these regions that should have none . if the shadow mask 36 does not perturb the x - ray flux significantly , the signal in these shadow regions can be taken as an estimate of the residual scatter detected at these points . this method has been previously employed for removing scatter from conventional projection radiographs . in contrast to the aforementioned method , which requires two scans — one with and without the shadow mask 36 , or two with the shadow mask 36 in different positions — it is possible to leave the shadow mask 36 in place during the entire scan , which has three desirable consequences . first , the x - ray flux remains unchanged between the scatter measurement and the measurement from which the scatter - free projection is estimated from since they are performed simultaneously . this feature simplifies the correction procedure . second , since the shadow mask 36 need not be moved , the overall design is simplified . finally , since it is not necessary to expose the patient twice to obtain both the diagnostic information and a measure of the scattered radiation , the overall radiation dose to the patient is reduced . there are a number of ways shadow masks can be designed to allow this , one example is shown in fig8 . in this case , the absorbing components of the mask 36 , such as an absorbing frame 37 and absorbing rods 39 , are slightly outside the desired field - of - view 46 ( i . e ., outside a clear , imaging aperture 41 ) but since scatter varies smoothly , the detector 30 “ sees ” approximately the same scatter at these points as it does within the desired region . referring to fig9 through 13 , an outline of the procedure for scatter estimation and correction is given as follows : 1 . with the shadow mask 36 between the source assembly 18 and the detector assembly 16 and no object in place , a “ blank scan ” is taken to determine where the shadows of the mask 36 lie on the detector 30 . 2 . this data is then processed to make two digital masks . the first defines regions - of - interest ( rois ) slightly smaller than the actual shadows . these rois will be used to estimate the residual scatter ( fig9 ). the second defines regions larger than the shadows ( fig1 ). these regions will be used in the correction process for the diagnostic data ( patient scan ). 3 . the patient is placed in the tomograph system 10 and scanned with the shadow mask 36 in place ( fig1 ). 4 . using the first mask , the scattered radiation is estimated by the values at each view angle or frame of the patient scan by the values in each frame at every non - zero mask position . 5 . the scatter at zero mask positions is then estimated using an interpolation scheme ( fig1 ). a variety of schemes can be used . 6 . the scattered information is either explicitly or implicitly ( e . g ., during the 3d image reconstruction process ) subtracted from the patient scan . using the second mask , information in the shadow regions of the patient scan is “ restored ” ( i . e ., as if there were no shadow mask present ). the preferred method for restoring this information is to use a penalized , weighted least - squares or penalized maximum likelihood 3d image reconstruction where the mask determines regions of missing data . alternatively , the information can be estimated on a frame - by - frame basis ( shown in fig1 ), which may , in some cases , be preferable because the amount of computation is significantly reduced . as noted , there are not only a variety of masks that can be employed but also a variety of estimation / restoration / reconstruction scenarios that can be used . detectors always have small spatial non - uniformities of response . sometimes these can be calibrated out by acquiring a sequence of images with the x - ray source off in order to obtain the “ dark signal ” or pedestal and another set with the x - ray source on with no object in place (“ blank scan ”) for determining the response or gain of each pixel . often , however , the signal from a pixel may vary unpredictably when scanning the object . if not taken into account , these signals tend to add coherently in a conventional fan - or cone - beam scanning geometry and will generate rings of higher or lower density in reconstructed images . the magnitude of the ring depends both on the magnitude of the detector defect and its spatial sharpness : sharp - edged defects will generate much larger artifacts than smooth - edged defects . a way to “ smooth ” any detector defect is to move the detector 30 along its track 35 a small amount randomly between each frame . this would typically be accomplished by moving the detector 30 under motor control between frames a random — but known — amount corresponding to a maximum of roughly ± 20 pixels . the data acquired in such a fashion are then shifted in the computer 12 after each image frame is collected such that it appears as if the detector 30 as in a single position . due to the random detector movement , non - uniformities that previously added coherently to produce disturbing visual distortions in the reconstructed images now do not significantly affect reconstructions . in cone - beam scanning , the more complete orbits described above ( i . e ., helical and wobbled ) further reduce the effect of isolated detector artifacts . with the area detector 30 , the pedestal or “ dark signal ” can vary considerably during the scan due to sensitivity of the readout electronics , which are proximal to the detector panel , to small variations in temperature . in ordinary radiography , these pedestal changes present little problem . in computed tomography , however , the changes may be larger than the signal of interest and it is therefore important that they be estimated throughout the scanning interval so that their effects can be removed either during or after the scanning . the varying pedestals in these devices are largely due to the readout amplifiers for each channel . in a device currently in use , each amplifier reads out a column of 256 channels ( pixels ) with half placed at the bottom edge and the remainder at the top edge of the detector ( the detector has an array of 512 × 512 pixels ). if the temperature of the readout chip changes , the offsets are changed for every pixel the amplifier reads out ( e . g ., an entire column ). referring now to fig1 a - 14 c , the detector 30 is shown with gate driver asics 61 and readout asics 63 . in order to estimate pedestals during scanning without acquiring frames when the x - ray source 26 is off ( which extends the overall scanning time ), portions of the detector 30 adjacent to the integrated circuits containing the readout amplifiers are covered with an x - ray absorbing material 60 ( i . e ., fig1 c ). the material 60 covers positions of a scintillator layer 62 . alternatively , the photodiodes on the flat - panel 30 can be masked from light . although this reduces the active area of the detector 30 slightly , the change in pedestal during the scan for each readout channel can be estimated from the change in the signal observed for the masked channels . there are numerous methods for estimating the pedestal of each channel from the masked - channel information ; the simplest assumes that the change in pedestal for each readout channel is the same for each detector element using a particular readout amplifier ( i . e ., a row or column of the array depending on orientation ) and corrects each digitized detector value by the same digital pedestal change . an alternative means for accomplishing this without detector masking — assuming the electronics of the detector 30 can be modified — is to have one or several electrical channels on each chip containing readout amplifiers not connected to the array ( i . e ., everything is similar to active channels but there is no possibility of a signal ). these dark channels can then serve as estimates for the change in pedestal for each readout chip during scanning . once the 3d volume of attenuation coefficients have been reconstructed from the projection data , the information can be reprojected so that it emulates that acquired using other , more conventional devices such as panoramic x - ray machines or conventional linear tomographs used in dentomaxillofacial imaging . this is useful both for comparing data acquired with the system 10 to previous patient scans and for providing a simple method for “ volume rendering ” or summarizing the 3d volumetric data in a format familiar to the clinician . generating the scans corresponding to those acquired using more conventional devices is summarized by the following procedure : 1 . 3d patient data is acquired and reconstructed using the system 10 . the result is a volume of linear attenuation coefficients at a particular energy . 2 . a linear system model describing how a panoramic x - ray machine or linear tomography generates images from an object is used to map the reconstructed image volume into projections that would have been acquired using a conventional system . this provides the sum of linear attenuation along source - detector paths collected in the conventional instrument . 3 . this information is multiplied by − 1 and exponentiated to provide the x - ray attenuation the conventional system would have observed . the information can then be displayed digitally or printed on paper or film for presentation . the key to the above procedure is the system model describing the imaging characteristics ( geometry , etc .) for each panoramic system or linear tomography . these linear models can be constructed using methods well known to developers of computed tomography software . a system constructed in accordance with the present invention may include one or more of the following : a vertically oriented cone - beam imaging system 10 capable of motions that allow a complete set of tomographic data to be acquired . this is accomplished by not only rotating the gantry 20 but by also allowing it to either pivot on an axis 53 orthogonal to the main rotation axis 51 or by moving the source assembly 18 and detector assembly 16 vertically , parallel to the main axis 51 of rotation during the scan such that the cone - beam completeness conditions are satisfied . a tomographic system 10 capable of acquiring a complete ct scan with high and isotropic spatial resolution for the dentomaxillofacial complex that additionally is capable of performing both single - and dual - kv (“ dual energy ”) imaging . the proposed scatter elimination technique using a coarse , high - transparency grid 32 ( or crossed coarse grids ) as opposed to fine granularity , relatively low - transparency grids . the shadow mask 36 containing both absorbing elements for scatter estimation and spatially varying filtration for dual - energy imaging for the purpose of correcting beam - hardening artifacts . the use of a shadow mask 36 that can be left in place during the diagnostic scan and methods for estimating the residual scatter and reconstructing the resulting data . the use of a low current scout scan to determine the appropriate detector offset in addition to estimating appropriate exposure time ( appropriate exposures are presently estimated using a scout scan ). the use of a detector that is smaller than necessary in conjunction with apparatus for moving the detector 30 relative to the source 26 and an appropriate x - ray collimator 28 to improve the precision of measurements by reducing scatter and by measuring some portions of the object twice relative to others . adapting the measurement system ( x - ray collimation and detector position ) dynamically to the size of the object to accomplish the above objectives ( i . e ., appropriate detector position for each frame is not estimated from scout scan but rather from previous frames ). estimating the pedestal ( or dark signal ) for each channel that does not require turning off the x - ray source 26 or acquiring additional “ blank ” frames between x - ray exposures . “ dithering ” the position of the detector 30 in order to reduce visually disturbing artifacts . generating “ panoramic ” or conventional linear tomography images from the reconstructed ct volume . extensive simulations and a lab - bench prototype model have been constructed and used for imaging phantoms and cadaver heads in order to assess the utility of the system 10 in treatment planning for dental implants . the spatial resolution of the system 10 is approximately 400 microns fwhm in all three spatial dimensions . moreover , complete data is acquired and both single - and dual - energy scanning have been accomplished . this is in contrast to conventional spiral ct where spatial resolution is closer to 1 mm in the transaxial direction and 1 . 5 - 2 mm in the axial direction . four tomographic slices of a human head are shown in fig1 a - 15 d . it is certainly possible to construct a small x - ray ct scanner for imaging the head , neck and extremities in a variety of ways . for example , one can attach a large enough amorphous silicon detector to a rotating gantry and use no additional scatter correction hardware or software and obtain images that may be adequate for planning dental implants in many cases . nevertheless , when it is desirable to be more quantitative to accurately assess bone quality , for instance , many of the features described above such as dual - energy imaging , scatter correction , and pedestal estimation methods will prove important . it is also possible to use detectors other than amorphous silicon imaging arrays . for example , image intensifiers and ccd cameras can be and have been used ( tomcat ). nevertheless , simple devices based on these detectors still suffer problems such as scatter , detector size issues , tomographic completeness problems , etc . the methods described above are solutions to those common problems and most do not depend on the type of detector used . while embodiments of the invention have been illustrated and described , it is not intended that these embodiments illustrate and describe all possible forms of the invention . rather , the words used in the specification are words of description rather than limitation , and it is understood that various changes may be made without departing from the spirit and scope of the invention .
a high spatial resolution x - ray computed tomography system is provided . the system includes a support structure including a gantry mounted to rotate about a vertical axis of rotation . the system further includes a first assembly including an x - ray source mounted on the gantry to rotate therewith for generating a cone x - ray beam and a second assembly including a planar x - ray detector mounted on the gantry to rotate therewith . the detector is spaced from the source on the gantry for enabling a human or other animal body part to be interposed therebetween so as to be scanned by the x - ray beam to obtain a complete ct scan and generating output data representative thereof . the output data is a two - dimensional electronic representation of an area of the detector on which an x - ray beam impinges . a data processor processes the output data to obtain an image of the body part .