Patent Publication Number: US-6215846-B1

Title: Densitometry adapter for compact x-ray fluoroscopy machine

Description:
CROSS-REFERENCE TO RELATED APPLICATIONS 
     This application is based on U.S. provisional application Ser. No. 60/080,164 filed Mar. 31, 1998 and is a continuation in part of U.S. application Ser. No. 09/006,358 filed Jan. 13, 1998 which is a continuation-in -part of PCT Application 97/02770 designating the United States filed Feb. 21, 1997 claiming the benefit of provisional application Ser. No. 60/011,993 filed Feb. 21, 1996. This provisional application is incorporated by reference herein. 
    
    
     STATEMENT REGARDING FEDERALLY SPONSORED RESEARCH OR DEVELOPMENT 
     FIELD OF THE INVENTION 
     The invention relates generally to x-ray equipment and in particular to an adapter for bone density measurements as may be used with compact x-ray fluoroscopy equipment used for orthopedic and similar procedures. 
     BACKGROUND OF THE INVENTION 
     Portable x-ray fluoroscopy machines provide an x-ray source held in opposition to an electronic image detector, typically on a C-arm, so that x-rays from the x-ray source are received by the image detector. The C-arm may slide through a collar so as to allow it to be rotated to different angles about the patient. Further, the collar may be supported by a pivoting arm providing additional freedom in the positioning of the C-arm. 
     When the C-arm is correctly positioned, the x-ray source is activated and x-rays pass through the patient to be received by the image detector which provides electronic signals to a video monitor. For larger mobile C-arm systems, the video monitor is typically held on a separate cart or may be suspended from the ceiling on a fixed bracket to be connected to the mobile unit when the mobile unit is in place. 
     With improvements in electronic hardware and in particular the development of compact image intensifiers and CCD video cameras, it has become possible to build extremely compact mobile C-arm systems. Such systems may make use of increasingly powerful desktop computer technology for image processing and other tasks and may use compact digital printers for producing images. 
     BRIEF SUMMARY OF THE INVENTION 
     The present invention provides an adapter that may convert a compact fluoroscopy machine or other mobile x-ray source into a precision quantitative densitometer suitable for measuring bone mass or density such as may be helpful in the treatment and detection of osteoporosis. 
     The invention provides a stand to be used with a fluoroscopy machine, the stand having a cradle for accurately locating the x-ray source and/or detector with respect to either the patient&#39;s forearm or foot. Special software is loaded to the computer of the fluoroscopy machine to operate the fluoroscopy machine in a quantitative dual energy mode and to adapt the fluoroscopy data to densiometric data. For fluoroscopy equipment not providing for digital imaging or that may not be easily operated in a dual energy mode, the invention includes a provision for a separate digital dual energy detector and if necessary an associated processing computer. 
     Other objects, advantages, and features of the present invention will become apparent from the following specification when taken in conjunction with the accompanying drawings. 
    
    
     BRIEF DESCRIPTION OF THE SEVERAL VIEWS OF THE DRAWINGS 
     FIG. 1 is a perspective view of the fluoroscopy machine suitable for use with the present invention showing a C-arm supporting an image intensifier/video camera and x-ray tube in opposition for rotation in a vertical plane, the C-arm held along a mid-line of a cart by an articulated arm attached to the side of the cart; 
     FIG. 2 is a side view in elevation of the cart of FIG. 1 showing a slide attaching the articulated arm to the side of the cart and showing a four-bar linkage motion of the arm for elevation of the C-arm; 
     FIG. 3 is a top view of the C-arm system of FIG. 1 with the articulated arm in partial phantom showing the four-bar linkage of the arm for extending the C-arm toward and away from the cart; 
     FIG. 4 is a detail fragmentary view of an outer pivot of the articulated arm attached to the C-arm such as allows limited pivoting of a plane of rotation of the C-arm about a vertical axis; 
     FIG. 5 is a detail view of the C-arm of FIG.  1  and the attached x-ray tube assembly showing the electrical cabling providing power to an x-ray tube power supply fitting into a groove in the C-arm and showing an abutment of the anode of the x-ray tube against the metal casting of the C-arm for heat sinking purposes; 
     FIG. 6 is a schematic block diagram of the fluoroscopy machine of FIG. 1 showing the path of control of a remote x-ray tube power supply by a microprocessor and the receipt of data from the image intensifier/video camera by the microprocessor for image processing; 
     FIGS. 7 and 8 are amplified images such as may be obtained by the system of FIG. 1 showing portions of the image having moving elements and portions having stationary elements; 
     FIG. 9 is a flow chart of a method of the present invention providing differently weighted noise reduction to different areas of the image based on motion in the areas of the image; 
     FIG. 10 is a figure similar to that of FIG. 7 showing an image of a rectilinear grid as affected by pincushion distortion in the image intensifier and video camera optics such as may provide a confusing image of a surgical tool being manipulated in real-time; 
     FIG. 11 is a figure similar to FIG. 10 showing equal areas of the image that encompass different areas of the imaged object, such variation as may affect quantitative bone density readings; 
     FIG. 12 is a plot of raw image data from the image intensifier/video camera as is translated into pixel brightness in the images of FIGS. 7,  8 ,  10 , and  11  by the microprocessor of FIG. 6 according to a non-linear mapping process such as provides noise equilibrium in the images and maximum dynamic range for clinical data; 
     FIG. 13 is a histogram plotting values of data from the image intensifier/video camera versus the frequency of occurrence of data values showing an isolated Gaussian distribution at the right most side representing unattenuated x-ray values; 
     FIG. 14 is a flowchart describing the steps taken by the programmed microprocessor of FIG. 6 to identify background pixels and remove them from a calculation of exposure rate used for controlling the remote x-ray tube power supply of FIG. 6; 
     FIG. 15 is a detailed block diagram of the first block of the flow chart of FIG. 14; 
     FIG. 16 is a first embodiment of the second block of the flow chart of FIG. 14; 
     FIG. 17 is a second embodiment of the second block of the flow chart of FIG. 14; 
     FIG. 18 is a detailed flow chart of the third block of the flow chart of FIG. 14; 
     FIG. 19 is a schematic representation of a distorted image of FIG. 11 and a schematic representation of a corresponding undistorted image showing the variables used in the mathematical transformation of the distorted image to correct for rotation and distortion; 
     FIG. 20 is a flow chart of the steps performed by the computer in correcting and transforming the image of FIGS. 11 and 19; 
     FIG. 21 is a perspective view of an occluder placed in an x-ray beam prior to an imaged object and used for calculating scatter; 
     FIG. 22 is a flow chart of the steps of calculating and removing scatter using the occluder of FIG. 21; 
     FIG. 23 is a cross-sectional view through the occluder of an imaged object of FIG. 21 along line  23 — 23 , aligned with a graph depicting attenuation of x-rays as a function distance along the line of cross-section as well as theoretical attenuation without scatter and scatter components; 
     FIG. 24 is a graphical representation of an adjustment of calculated scatter from the image of FIG. 23 based on normalizing points established by the occluder of FIG. 21; 
     FIG. 25 is a perspective view similar to that of FIG. 1 showing a C-arm system similar to that of FIG. 1 in position on the densitometry cradle of the present invention to provide a beam of x-rays along a horizontal axis across the top of the cradle; 
     FIG. 26 is a fragmentary exploded view of the x-ray detector and x-ray source of the C-arm system of FIG. 25 removed from the cradle and showing a removable foot positioner also removed from the cradle; 
     FIG. 27 is an enlarged, fragmentary view of FIG. 26 showing the alternative fitting of a forearm positioner or the foot positioner within a channel of the cradle along the path of x-rays between the x-ray source and x-ray detector; 
     FIG. 28 is a top plan view of the cradle of FIG. 25 but with the C-arm removed, showing the forearm positioner in use with a patient&#39;s arm; 
     FIG. 29 is a side elevational view of the cradle of FIG. 28; 
     FIG. 30 is a top plan view of the cradle of FIG. 28 but with the foot positioner in use with a patient&#39;s leg and showing the use of an auxiliary pancake image intensifier for use with fluoroscopy or other x-ray equipment not having suitable dual energy or digital imaging capabilities; 
     FIG. 31 is a cross sectional view of the pancake image intensifier of FIG. 30 with the protective shrouding removed and taken along a plane including the x-ray beam axis showing the compact, high distortion configuration and an attached processing computer; 
     FIG. 32 is a flow chart of software executed by a processing computer associated with the -x-ray detector for providing quantitative densiometric data; 
     FIG. 33 is a block diagram of the steps of scatter correction of the present invention; and 
     FIG. 34 is a block diagram illustrating the removal of line correlated noise per the present invention. 
    
    
     DETAILED DESCRIPTION OF THE INVENTION 
     C-arm Support Mechanism 
     Referring now to FIG. 1, an x-ray machine  10  per the present invention includes a generally box-shaped cart  12  having castors  14  extending downward from its four lower corners. The castors  14  have wheels rotating about a generally horizontal axis, and swiveling about a generally vertical axis passing along the edges of the cart  12 . Castors  14 , as are understood in the art, may be locked against swiveling and/or against rotation. 
     With one castor  14  locked and the others free to rotate and swivel, a pivot point  15  for the cart  12  is established with respect to the floor such as may be used as a first positioning axis  11  for the x-ray machine  10 . 
     Positioned on the top of the cart  12  is a turntable  16  holding a video monitor  18  and attached keyboard  20  for swiveling about a vertical axis for convenience of the user. The video monitor  18  and the keyboard  20  may swivel separately so that one operator may view the video monitor  18  while a second operates the keyboard  20 . 
     The video monitor  18  and the keyboard  20  allow for control of a computer  22  contained in a shelf on the cart  12  open from the front of the cart  12 . The computer  22  may include a general microprocessor-type processor  23  and a specialized image processor  27  for particular functions as will be described. The computer  22  further includes a number of interface boards allowing it to provide control signals to various components of the x-ray machine  10  as will be described and to receive x-ray image data. In addition, the computer  22  receives signals from a foot switch  61  that is used to activate the x-ray system for a brief exposure. Control of the computer  22  may also be accomplished through a remote control wand  63  of a type known in the art. 
     Referring now also to FIG. 2, attached to the right side of the cart  12  is a horizontal slide  24  positioned to provide an attachment point  26  for an articulated arm  19  supporting a substantially circular C-arm  56 , which in turn holds an x-ray tube  68  and an image intensifier  82  and camera  84 , in opposition, and facing each other as will be described below. The function of the x-ray tube, the image intensifier and the camera are well known in the prior art in the use of mobile C-arm type x-ray devices used for image display and are described in U.S. Pat. No. 4,797,907 hereby incorporated by reference as part of the prior art. The C-arm may be mass balanced, that is to say its weight may be distributed to reduce its tendency to rotate through collar  54  so that minimal frictional pressure may be used to prevent it from moving. 
     The articulated arm  19  may be slid horizontally toward the front of the cart  12  to provide a second positioning axis  25  of the x-ray machine  10 . A first pulley  28  is rotatively fixed in a vertical plane, attached to the portion of the slide  24  that may move with respect to the cart  12 , and is pivotally attached to a rigid arm  30  extending toward the front of the cart  12 . The other end of the rigid arm  30  supporting a second pulley  32  is also mounted to swivel with respect to arm  30 . A belt  34  wraps around a portion of the circumference of each of pulleys  28  and  32  and is affixed at one point along that circumference to each of the pulleys  28  and  32  so that pivoting motion of the arm  30  about the center point  26  of pulley  28  causes rotation of pulley  32  so that it maintains a fixed rotational orientation with respect to the cart  12  as pulley  32  and hence C-arm  56  is moved up and down along a third axis  37 . The linkage, so created, is a variation of the “four bar linkage” well known in the art. 
     Helical tension springs (not shown for clarity) balance the pulley  32  in rotative equilibrium about point  26  against the weight of the articulated arm  19 , C-arm  56 , and other devices attached to the arm  19 . 
     Attached to pulley  32  is a third pulley  36  extending in a generally horizontal plane perpendicular to the plane of pulley  32 . The third pulley  32  is attached pivotally to a second rigid arm  40  which at its other end holds another pulley  38  positioned approximately at the midline  41  of the cart  12 . The midline  41  symmetrically divides the left and right sides of the cart  12 . 
     Portions of the circumference of pulleys  36  and  38  are also connected together by a belt  44  so as to form a second four bar linkage allowing pulley  38  to move toward and away from the cart  12 , along a fourth positioning axis  45 , with pulley  38  and C-arm  56  maintaining their rotational orientation with respect to cart  12 . 
     Referring now to FIG. 4, pulley  38  includes a center shaft member  50  having a coaxial outer collar  52  to which belt  44  is attached. A stop  55  attached to the shaft  50  limits the motion of the collar  52  in rotation with respect to the shaft  50  to approximately 26 degrees. Frictional forces between shaft  50  and collar  52  cause shaft  50  to maintain its rotational orientation with respect to collar  52  and hence with respect to pulley  36  until sufficient force is exerted on shaft  50  to displace it with respect to collar  52 . Thus pressure on the C-arm  56  can provide some pivoting motion of the C-arm about the axis of the pulley along the fifth positional axis  55 . 
     Referring now to FIGS. 1,  3  and  4 , attached to the shaft  50  is a C-arm collar  52  supporting the arcuate C-arm  56  curving through an approximately 180 degree arc in a vertical plane substantially aligned with the midline  41  of the cart  12  as has been mentioned. The shaft  50  may connect to collar  52  so that the latter may swivel in about a horizontal axis bisecting the circle of the C-arm  56 . This axis may be aligned with the center of mass of the C-arm  56  so that there is not a self-righting tendency of the C-arm or the axis may be placed above the axis of the C-arm so as to provide for a beneficial self righting action. This motion is orthogonal to that provided by motion of shaft  50  and may augment that provided by the castors  14 . Techniques of balancing the C-arm in its various rotational modes, when this is desired, is taught by U.S. Pat. No. 5.038,371 to Janssen issued Aug. 6th, 1991 and hereby incorporated by reference as exemplifying the known prior art understood to all those of ordinary skill in the art. 
     As described above, motion of the collar  52  may be had in a vertical manner by means of the parallelogram linkage formed by pulleys  28  and  32  of the articulated arm  19  as shown in FIG.  2 . Forward and backward motion away from and toward the cart  12  may be had by the second four bar linkage formed from pulleys  36  and  38 . A slight pivoting of the C-arm  56  about a vertical axis slightly to the rear of the collar  52  and concentric with the axis of pulley  38  may be had by means of the rotation between collar  52  and  50  of FIG.  4 . Greater rotation of the C-arm about the vertical axis passing through pivot point  15  may be had by rotation of the cart about one of its stationary castors  14 . Thus, considerable flexibility in positioning the C-arm may be had. 
     Referring now to FIG. 5, the C-arm  56  is an aluminum casting having formed along its outer circumference a channel  58  into which a cable  60  may be run as will be described. C-arm  56  has a generally rectangular cross-section taken along a line of radius of the C-arm arc. Each corner of that rectangular cross-section holds a hardened steel wire  62  to provide a contact point for corner bearings  64  within the collar  52 . The corner bearings  64  support the C-arm  56  but allow movement of the C-arm  56  along its arc through the collar  52 . 
     A cable guide pulley  66  positioned over the channel  58  and having a concave circumference feeds the cable  60  into the channel  58  as the C-arm moves preventing tangling of the cable  60  or its exposure at the upper edge of the C-arm  56  when the C-arm  56  is rotated. The excess length of cable  60  loops out beneath the collar  52 . 
     X-Ray Tube Cooling 
     Referring now to FIGS. 5 and 6, the C-arm supports at one end a generally cylindrical x-ray tube  68  having a cathode  70  emitting a stream of electrons against a fixed anode  72 . The conversion efficiencies of x-ray tubes are such that the anode  72  can become quite hot and typically requires cooling. In the present invention, the anode  72  is positioned to be bolted against the aluminum casting of the C-arm  56  thereby dissipating its heat into a large conductive metal structure of the C-arm  56 . 
     The x-ray tube  68  is connected to an x-ray tube power supply  74  which separately controls the current and voltage to the x-ray tube  68  based on signals received from the computer  22  as will be described. The control signals to the x-ray tube power supply  74  are encoded on a fiber optic within the cable  60  to be noise immune. Low voltage conductors are also contained within cable  60  to provide power to the x-ray tube power supply  74  from a low voltage power supply  76  positioned on the cart  12 . 
     During operation, an x-ray beam  80  emitted from the x-ray tube  68  passes through a patient (not shown) and is received by an image intensifier  82  and recorded by a charge couple device (“CCD”) camera  84  such as is well known in the art. The camera provides digital radiation values to the computer  22  inversely proportional to the x-ray absorption of the imaged object for processing as will be described below. Each radiation value is dependent on the intensity of x-ray radiation received at a specific point on the imaging surface of the image intensifier  82 . 
     Image Noise Reduction 
     Referring now to FIGS. 6 and 7, the data collected by the CCD camera  84  may be used to provide an image  86  displayed on video monitor  18 . As will be described in more detail below, the CCD camera receiving a light image from the image intensifier  82  at a variety of points, provides data to the computer which maps the data from the CCD camera  84  to a pixel  88  in the image  86 . For convenience, the data from the CCD camera  84  will also be termed radiation data reflecting the fact that there is not necessarily a one-to-one correspondence between data detected by the CCD camera  84  and pixels  88  displayed on the video monitor  18 . 
     The CCD camera  84  provides a complete set of radiation data for an entire image  86  (a frame) periodically once every “frame interval” so that real-time image of a patient placed within the x-ray beam  80  may be obtained. Typical frame rates are in the order of thirty frames per second or thirty complete readouts of the CCD detector area to the computer  22  each second. 
     Each frame of data is stored in the memory of the computer  22  and held until after complete storage of the next frame of data. The memory of the computer  22  also holds an average frame of data which represents an historical averaging of frames of data as will now be described and which is normally used to generate the image on the video monitor  18 . 
     In a typical image  86 , there will be some stationary object  90  such as bone and some moving object  92  such as a surgical instrument such as a catheter. In a second image  86 ′ taken one frame after the image  86 , the bone  90  remains in the same place relative to the edge of the image  86  and  86 ′, however the surgical instrument  92  has moved. Accordingly, some pixels  88 ′ show no appreciable change between images  86  and  86 ′, whereas some other pixels  88 ″ show a significant change between images  86  and images  86 ′. 
     Referring now to FIG. 9, as data arrives at the computer  22 , the computer  22  executes a stored program to compare current pixels of the image  86 ′ to the last pixels obtained from image  86  as indicated by process block  94 . This comparison is on a pixel by pixel basis with only corresponding pixels in the images  86  and  86 ′ compared. The difference between the values of the pixels  88 , reflecting a difference in the amount of x-ray flux received at the CCD camera  84 , is mapped to a weight between zero and one, with greater difference between pixels  88  in these two images corresponding to larger values of this weight w. This mapping to the weighting is shown at process block  96 . 
     Thus pixels  88 ″, whose value changes almost by the entire range of pixel values between images  86  and  86 ′, receive a weighting of “one” whereas pixels  88 ′ which have no change between images  86  and  86 ′ receive a value of zero. The majority of pixels  88  being neither unchanged nor radically changed will receive a value somewhere between zero and one. 
     Generally, because the amount of x-ray fluence in the beam  80  is maintained at a low level to reduce the dose to the patient, the images  86  and  86 ′ will have appreciable noise represented as a speckling in the images  86  and  86 ′. This noise, being of random character, may be reduced by averaging data for each pixel  88  over a number of frames of acquisition effectively increasing the amount of x-ray contributing to the image of that pixel. 
     Nevertheless, this averaging process tends to obscure motion such as exhibited by the surgical instrument  92 . Accordingly, the present invention develops an average image combining the values of the pixels acquired in each frame  86 ,  86 ′ in which those pixels in the current image  86 ′ which exhibit very little change between images  86  and  86 ′ contribute equally to the average image, but those pixels in the current image  86 ′ that exhibit a great degree of change between images  86  and  86 ′ are given a substantially greater weight in the average image. In this process, a compromise is reached between using historical data to reduce noise and using current data so that the image accurately reflects changes. Specifically, the value of each pixel displayed in the image is computed as follows. 
     
       
           P   i =(1− w ) P   i−1   +wP   i,t   (1)  
       
     
     where P i−1  is a pixel in the previous average image, w is the weighting factor described above and P i,t  is the current data obtained from the CCD camera  84 . This effective merger of the new data and the old data keyed to the change in the data is shown at process block  98 . 
     Image Intensifier Distortion 
     Referring now to FIG. 10, an image  86 ″ of a rectilinear grid  100  positioned in the x-ray beam  80  will appear to have a barrel or pincushion shape caused by distortion of the image intensifier  82  and the optics of the CCD camera  84 . During a real-time use of the image  86 ″ by a physician, this distortion may cause confusion by the physician controlling a tool  102 . For example, tool  102  may be a straight wire shown by the dotted line, but may display an image  86 ′ as a curved wire whose curvature changes depending on the position of the tool  102  within the image  86 . This distortion thus may provide an obstacle to a physician attempting to accurately place the tool  102  with respect to an object within the image  86 ′. 
     Referring now to FIG. 11, the distortion of image  86 ″ also means that two equal area regions of interest  105  (equal in area with respect to the image) do not encompass equal areas of the x-ray beam  80  received by the image intensifier  82 . Accordingly, if the data from the CCD camera  84  is used for quantitative purposes, for example to deduce bone density, this distortion will cause an erroneous variation in bone density unrelated to the object being measured. 
     Accordingly, the present inventors have adopted a real-time digital re-mapping of radiation data from the CCD camera  84  to the image  86  to correct for any pincushion-type distortion. This remapping requires the imaging of the rectilinear grid  100  and an interpolation of the position of the radiation data received from the CCD camera  84  to new locations on the image  86 ″ according to that test image. By using digital processing techniques in a dedicated image processor  27 , this remapping may be done on a real-time basis with good accuracy. 
     Referring to FIG. 19, there are two types of distortion, isotropic and anisotropic. Isotropic distortion is rotationally symmetric (e.g. like barrel and pin cushion distortion). Anisotropic distortion is not rotationally symmetric. Both types of distortion and rotation are so-called third order aberrations which can be written in the form: 
     
       
           Dx=r   2 ( Du−dv )  (2) 
       
     
     
       
           Dy=r   2 ( Dv+du )  (3) 
       
     
     where Dx and Dy are pixel shifts due to distortion; r is the distance from the correct position to the optical axis and D and d are distortion coefficients which are constant and u and v are correct pixel positions. 
     Referring also FIG. 2, received image  86  may exhibit pin cushion distortion evident if an image  86  of the rectilinear grid  100  is made. The distortion is caused by the pixel shifts described above. 
     Equations 1 and 2 may be rewritten as third order two-dimensional polynomials, the case for equation (1) following: 
     
       
           x= ( a   x   +e   x   v+i   x   v   2   +m   x   v   3 )+( b   x   +f   x   v+j   x   v   2   +n   x   v   3 ) u +( c   x   +g   x   v+k   x   v   2   +o   x   v   3 ) u   2 +( d   x   +h   x   v+l   x   v   2   +p   x   v   3 ) u   3   (4) 
       
     
     In these polynomials, a x  and a y  govern the x and y translation of the image, e x  and b y  take care of scaling the output image, while e y  and b x  enable the output image to rotate. The remaining higher order terms generate perspective, sheer and higher order distortion transformations as will be understood to those of ordinary skill in the art. Thirty-two parameters are required for the two, third order polynomials. These parameters may be extracted by a program executed by the computer in an off-line (non-imaging) mode after imaging the known grid  100  and comparing the distorted image of the grid  100  to the known grid  100  to deduce the degrees of distortion. 
     Referring now to FIG. 19 in a first step of the correction process, the grid  100  is imaged as indicated by process block  160  to determine the exact type of distortion present and to obtain values for the coefficients a through p of the above referenced polynomial equations. 
     At process block  166 , these parameters may be input to the computer  22  and used at a transformation of received image  86  into image data  164  as indicated by process block  168 . For rotation of the image  164 , new parameters of the polynomials may be entered by means of hand-held remote control wand  63  shown in FIG.  1 . 
     The transformation process generally requires a determination of the pixel shift for each radiation pixel  163  of the input image  86  which in turn requires an evaluation of the polynomials whose coefficients have been input. A number of techniques are known to evaluate such polynomials including a forward differencing technique or other techniques known in the art. These transformations provide values of u and v for an image pixel  170  corresponding to a particular radiation pixel  163 . 
     After the transformation of process block  168 , the u, v locations of the radiation pixels will not necessarily be centered at a pixel location defined by the hardware of the video monitor  18  which usually spaces pixels  170  at equal distances along a Cartesian axis. Accordingly, the transformed pixels must be interpolated to actual pixel locations as indicated by process block  172 . 
     A number of interpolation techniques are well known including bilateral and closest neighbor interpolation, however in the preferred embodiment, a high resolution cubic spline function is used. A given value of an interpolated pixel  170  (P int ) is deduced from a 4×4 block of transform pixels (P i,j ) in which it is centered as follows: 
     
       
         P int   =f ( n− 2) X   1   +f ( n− 1) X   2   +f ( n ) X   3   +f ( n+ 1) X   4   (5)  
       
     
     where: 
     
       
           X   i   =f ( m− 2) P   i,1   +f ( m− 1) P   i,2   +f ( m ) P   i,3   +f ( m+ 1) P   i,4   (6)  
       
     
     where: 
     
       
           f ( x )=( a+ 2) x   3 +−( a+ 3) x   2 +1 for xε[0,1]; 
       
     
     
       
           f ( x )= ax   3 +−5 ax   2 +8 ax− 4 a  for xε[1,2];  (7)  
       
     
     f(x) is symmetrical about zero. In the preferred embodiment a=−0.5 and where m and n are fractions indicating the displacement of the neighboring pixels P i,j  with respect to P int  in the x and y directions, respectively. 
     At process block  180 , the transformed and interpolated image is displayed. 
     Noise Equalization 
     Referring now to FIG. 12, the radiation data from the CCD camera  84  are mapped to the brightness of the pixels of the image  86  according to a second transformation. In the preferred embodiment, this mapping between CCD radiation data and image pixel brightness follows a nonlinear curve  103  based on the hyperbolic tangent and being asymptotically increasing to the maximum CCD pixel value. This curve is selected from a number of possibilities so that equally wide bands of image pixel brightness  104  and  106  have equal amounts of image noise. The curve  103  is further positioned to provide the maximum contrast between clinically significant tissues in the image. 
     Exposure Control 
     The noise in the image  86  is further reduced by controlling the fluence of the x-ray beam  80  as a function of the density of tissue of the patient within the beam  80 . This density is deduced from the image  86  itself produced by the CCD camera  84 . In response to the image data, a control signal is sent via the fiber optic strand within the cable  60  to the x-ray tube power supply  74  positioned adjacent to the x-ray tube  68  (shown in FIG.  5 ). By positioning the x-ray tube power supply  74  near the x-ray tube  68 , extremely rapid changes in the power supplied to the x-ray tube  68  may be obtained. Distributed capacitances along high tension cables connecting the x-ray tube  68  to a stationary x-ray tube power supply are thus avoided in favor of low voltage cable  60 , and the shielding and inflexibility problems with such high tension cables are also avoided. 
     Automatic Technique Control 
     Referring now to FIGS. 13 and 14, a determination of the proper control signal to send to the x-ray tube power supply  74  begins by analyzing the image data  86  as shown in process block  120 . The goal is to provide for proper exposure of an arbitrary object placed within the x-ray beam  80  even if it does not fill the field of view of the CCD camera  84 . For this reason, it is necessary to eliminate consideration of the data from the CCD camera  84  that form pixels in the image corresponding to x-rays that bypass the imaged object and are unattenuated (“background pixels”). These background pixels may be arbitrarily distributed in the image  86  and therefore, this identification process identifies these pixels based on their value. To do this, the computer  22  collects the values of the pixels from the CCD camera  84  in a histogram  122  where the pixels are binned according to their values to create a multiple peaked plot. The horizontal axis of the histogram  122  may for example be from 0 to 255 representing 8 bits of gray scale radiation data and the vertical axis may be a number of pixels having a particular value. 
     If there is a histogram value at horizontal value  255 , and the maximum gray scale exposure recorded, the entire area of the histogram  122  is assumed to represent the imaged object only (no background pixels). Such a situation represents an image of raw radiation only or a high dose image of a thin object with possible clipping. In assuming that the whole histogram  122  may be used to calculate technique without removal of background pixels, a reduced exposure rate will result as will be understood from the following description and the peak classification process, to now be described, is skipped. 
     Otherwise, if there are no pixels with the maximum value of  225 , the present invention identifies one peak  124  in the histogram  122  as background pixels indicated by process block  120  in FIG.  14 . In identifying this peak  124 , the computer  22  examines the histogram  122  from the brightest pixels (rightmost) to the darkest pixels (leftmost) assuming that the brightest pixels are more likely to be the unattenuated background pixels. The process block  120  uses several predetermined user settings as will be described below to correctly identify the peak  124 . 
     Once the peak  124  has been identified, the pixels associated with that peak are removed per process block  126  by thresholding or subtraction. In the thresholding process, pixels above a threshold value  138  below the peak  124  are considered to be background pixels and are omitted from an exposure rate calculation. In the subtraction method, the peak  124  itself is used as a template to identify pixels which will be removed. 
     At process block  128 , an exposure rate is calculated based on the values of the pixels in the remaining histogram data and at process block  130 , an amperage and voltage value are transmitted via the cable  60  to the x-ray tube power supply and used to change the power to the x-ray tube. Generally, if the exposure rate is above a predetermined value, the amperage and voltage are adjusted to cut the x-ray emission from the x-ray tube, whereas if the exposure rate is below the predetermined value, the amperage and voltage are adjusted to boost the exposure rate to the predetermined value. 
     Referring now to FIGS. 13,  14  and  15 , the process of identifying background pixels will be explained in more detail. Process block  120  includes as a first step, an identification of a right most peak  124  in the histogram  122  (shown in FIG. 13) as indicated by subprocess block  132 . 
     At succeeding subprocess block  134 , this right most peak  124  is compared against three empirically derived parameters indicated in the following Table 1: 
     
       
         
           
               
               
             
               
                 TABLE 1 
               
               
                   
               
             
            
               
                 Minimum Slope Range 
                 Minimum necessary pixel range for 
               
               
                 (MSR) 
                 which the slope of the peak must be 
               
               
                   
                 monitonically increasing. 
               
               
                 Histogram Noise Level 
                 Minimum height of the maximum 
               
               
                 (HNL) 
                 value of the peak. 
               
               
                 Maximum Raw Radiation Width 
                 Maximum width of the detected peak 
               
               
                 (MRRW) 
                 with respect to the width of the entire 
               
               
                   
                 histogram. 
               
               
                   
               
            
           
         
       
     
     Specifically at subprocess block  134 , each identified peak  124  is tested against the three parameters indicated in Table 1. In the description in Table 1, “width” refers to the horizontal axis of the histogram  122  and hence a range of pixel values, whereas “height” refers to a frequency of occurrence for pixels within that range, i.e., the vertical axis of the histogram  122 . 
     These first two tests, MSR and HNL, are intended to prevent noise peaks and peaks caused by bad imaging elements in the CCD camera  84  or quantization of the video signal in the A to D conversion from being interpreted as background pixels. 
     Peaks  124  with a suitable stretch of monotonically increasing slope  131  (shown in FIG. 13) according to the MSR value and that surpass the histogram noise level HNL  133  are evaluated against the MRRW parameter. This third evaluation compares the width  135  of the histogram  122  against the width of the entire histogram  122 . The MRRW value is intended to detect situations where the imaged object completely fills the imaging field and hence there are no unattenuated x-ray beams or background pixels being detected. A valid peak  124  will normally have a width  135  more than 33% of the total width of the histogram  122 . 
     At decision block  136  if the peak  124  passes the above tests, the program proceeds to process block  126  as indicated in FIG.  14 . Otherwise, the program branches back to process block  132  and the next peak to the left is examined against the tests of process block  134  until a passing peak is found or no peak is found. If no peak is found, it is assumed that there are no background pixels and a raw exposure value is calculated from all pixels as described above. 
     Assuming that a peak  124  passes the tests of Table 1, then at process block  126  background pixels identified by the peak  124  selected at process block  120  are eliminated. 
     In a first method of eliminating background pixels indicated at FIG. 16, a magnitude threshold  138  within the histogram  122  is identified. Pixels having values above this threshold will be ignored for the purpose of selecting an exposure technique. The threshold  138  is established by identifying the center  140  of the peak  124  (its maximum value) and subtracting from the value of the center a value s being the distance between the start of the peak  124  as one moves leftward and the maximum  140 . The area under the histogram  122  for values lower than the threshold  138  is computed to deduce a raw exposure value which will be used as described below. 
     In a second embodiment, the shape of the histogram peak  124  from the start of the peak as one moves leftward to its maximum  140  is reflected about a vertical line passing through the maximum  140  and subtracted from the histogram peak  124  to the left of the vertical line. This approach assumes that the peak  124  of the background pixels is symmetrical and thus this method better accommodates some overlap between the object pixels and the background pixels in the histogram  122 . Again, the remaining pixels of the histogram  122  are summed (by integration of the area under the histogram  122  minus the area of the peak  124  as generated by the reflection) to provide a raw exposure value. 
     Referring now to FIG. 18, the raw exposure value is transformed by the known transfer characteristics of the CCD camera (relating actual x-ray dose to pixel value) to produce a calculated current exposure rate as indicated at process block  144 . 
     Referring to process block  146 , the current exposure rate is next compared to a reference exposure rate, in the preferred embodiment being 1.0 mR per frame, however this value may be refined after further clinical testing. If at process block  148 , the current exposure rate is within a “half fine-tune range” of the reference exposure rate, then the program proceeds to process block  150 , a fine tuning process block, and the amperage provided to the x-ray tube are adjusted in accordance to the disparity between the amperage and reference exposure rate. That is, if the current exposure is greater than the reference exposure rate, the amperage to the x-ray tube is reduced. The new value of amperage is compared against a predetermined range of amperage values (maximum beam current and minimum beam current values) so that the amperage value may never vary outside of this range. 
     If at decision block  148 , the current exposure rate is outside of the half fine tune range established at decision block  148 , a more substantial adjustment process is undertaken. Generally, the exposure provided by an x-ray system will follow the following equation: 
     
       
           X≈s·mA·kVp   n .  (8)  
       
     
     where: 
     s is seconds, 
     mA is the amperage provided to the x-ray tube, 
     kVp is the voltage provided to the x-ray tube, and 
     n is a power factor dependent on the geometry of the machine and the particular kind of object being imaged. 
     Generally, the value of n will not be known in advance. Accordingly in the more substantial correction process, n is deduced by obtaining two different exposures for equal predetermined intervals with different kVp values so that the value of n may be deduced. 
     At decision block  152 , it is determined whether a first or second reference exposure is to be obtained. If the first reference exposure was just obtained, the program proceeds to process block  154  and a new value of kVp is determined for a second exposure. In this case, the first exposure used will be that which was employed to produce the histogram  122  as previously described. 
     If the comparison of process block  148  indicated that the exposure rate was too high, a lower kVp value is selected; and conversely, if the exposure at process block  148  indicated the exposure was too low, an increased value of kVp is provided. The new kVp value for the second exposure must be within a predetermined range of kVp values established by the user. Mathematically, the kVp value selected may be described as: 
     
       
           kVp   2   =kVp   1   +a ( dkVp )  (9) 
       
     
     where a is a step factor and 
     dkVp is a minimum practical change in tube voltage. 
     Two preferred means of selecting may be used: one providing linear and one providing logarithmic scaling. Such scaling techniques are well understood to those of ordinary skill in the art. 
     If at decision block  152 , a second frame has already been taken with the new voltage value, then the program proceeds to process block  156  and the value of n in equation (9) is calculated. If the value of amperage is held constant between the first and second frame, the value of n may be determined according to the following equation:              n   =     log            X   2       X   1       /   log            kVp   2       kVp   1                 (   10   )                         
     where X 1  and X 2  are the measured exposure rates at the first and second frames, respectively and 
     kVp 1  and kVp 2  are the two x-ray tube voltages during the first and second frames. 
     At process block  158 , this value of ‘n’ is checked against threshold values intended to detect whether an erroneous value of n has been produced as a result of ‘clipping’ in the radiation data used to calculate exposure. As is understood in the art, clipping occurs when an increased dose of an element of the CCD camera produces no increase in the camera&#39;s output. 
     At decision block  158 , if the value of n calculated at process block  156  is greater than or equal to one, it is assumed to be valid and the program proceeds to process block  160  where kVp and mA are adjusted by setting mA equal to a maximum reference value and calculating kVp according to the following equation:                kVp   new     =         kVp   2          (         X   ref          mA   2           X   2          mA   ref         )            1   /   n               (   11   )                         
     where kVp new  and mA new  are the settings for the next frame to be shot. 
     If the resulting kVp value conflicts with the minimum system, kVp, kVp is set to the minimum system value and mA is calculated according to the following equation using the mA and kVp value of the second frame.                mA   new     =       mA   2            X   ref       X   2              (       kVp   2       kVp   min       )     n               (   12   )                         
     If the value of n in decision block  158  is less than one, then at process block  162 , n is tested to see if it is less than zero. This value of n is realized when the exposure rate of the second frame changes in the opposite direction of the tube voltage. This suggests a clipped histogram and therefore the program branches back to process block  154  to obtain a new second frame. This condition may also arrive from object motion between the first and second frame. 
     On the other hand, if at decision block  162 , n is not less than zero (e.g. n is between zero and 1), the program proceeds to process block  166 . Here it is assumed that because the sensitivity of the exposure rate on change in kVp is low, there may be some partial clipping. New values of kVp and mA are then computed and used with the previous second frame values to calculate a new n as follows. Generally, if kVp and mA are high, they are both lowered and if kVp and mA are low, they are both raised. 
     Scatter Reduction 
     Referring now to FIG. 1, the image produced by the present invention may be used for quantitative analysis including, for example, that of making a bone density measurement. It is known to make bone density analyses from x-ray images through the use of dual energy techniques in which the voltage across the x-ray tube is changed or a filter is periodically placed within the x-ray beam to change the spectrum of the x-ray energy between two images. The two images may be mathematically processed to yield information about different basis materials within the image object (e.g. bone and soft tissue). For these quantitative measurements, it is desirable to eliminate the effect of scatter. 
     Referring now to FIG. 23 in imaging a patient&#39;s spine  200 , for example, x-rays  202  are directed from an x-ray source  201  through the patient  199  to pass through soft tissue  204  surrounding a spine  200 . Certain of the x-rays  202  are blocked by the spine  200  and others pass through the spine  200  to be recorded at the image intensifier  206 . An attenuation image  208  measured by an image intensifier measures those x-rays passing through the patient  109 . 
     A portion  210  of the attenuation image directly beneath the spine  200  records not only those x-rays  202  passing through the spine  200  and the soft tissue  204  above and below it, but also scattered x-rays  212  directed, for example, through soft tissue  204  to the side of the spine  200  but then scattered by the soft tissue to proceed at an angle to the portion  210  of the attenuation image  208  beneath the spine  200 . Because the scattered x-rays  212  do not carry information about the attenuation of the spine  200 , they are desirably removed from the image  208  prior to its use in quantitative measurement. 
     For this purpose, the present invention uses an occluder  214  being an x-ray transparent plate such as may be constructed of Plexiglas and incorporating into its body, a plurality of x-ray blocking lead pins  216 . Preferably these pins are placed so as to project images  218  onto the image  208  received by the image intensifier  206  in positions outside an image  220  of the spine  200 . Generally therefore, the pins  216  are placed at the periphery of the occluder  214 . The pins  216  are sized so as to substantially block all direct x-rays from passing through them but so that their images  218  include a significant portion of scattered x-rays  212 . 
     Referring now to FIG. 22 at a first step of a scatter reduction operation with the occluder  214  of FIG. 21, an image is acquired of the imaged object, for example, the spine  200  and its surrounding soft tissue  204  (not shown in FIG. 21) including the images  218  of the pins  216 . This acquisition is indicated by process block  221  of FIG.  22 . 
     The pins  216  are held in predetermined locations with respect to the image  208  so that their images  218  may be readily and automatically identified. Preferably the pins  216  are placed at the interstices of a Cartesian grid, however, other regular patterns may be chosen. The image  208  may be corrected for pincushion type distortion, as described above, so that the locations of the pins  216  may be readily located in the image based on their known positions in the occluder  214 . 
     At each pin image  218 , a value  222  indicating the magnitude of the received x-rays, shown in FIG. 23, may be ascertained. This value  222  measures the scatter received in the vicinity of image  218  caused generally by the effect of the soft tissue  204  and possible secondary scatter effects in the image intensifier  206 . Values  222  are recorded, as indicated by process block  224 , for each pin image  218 . From these values, a set of normalizing points are established. 
     The image  208  is then used to derive a scatter map. Referring to FIG. 23, generally the amount of scatter at a given point will be a function of how many x-ray photons are received at points adjacent to the given point. For example, comparing the image  208  to a theoretical scatterless image  228  generally in an attenuated region  230  of the image  208  (e.g., under the spine  200 ), scatter will increase the apparent value in the image  208  as a result of radiation from nearby low attenuation regions scattering into the high attenuation region  230 . Conversely the apparent value at a low attenuation region  232  will be decreased because of the scatter into the high attenuation region. 
     A map of the scattered radiation may thus be modeled by “blurring” the image  208 . This blurring can be accomplished by a low pass filtering of the image  208 , i.e., convolving the image  208  with a convolution kernel having rectangular dimensions corresponding to the desired low pass frequency cut off. The effect is an averaging of the image  208  producing scatter map  234 . 
     The image used to produce the scatter map  234  is an attenuation image  208  obtained from the patient  199  without the occluder  214  in place, or may be an image  208  including the images  218  of the pins  216  but with the latter images  218  removed based on knowledge of their location. This removal of images  218  may substitute values of the image  208  at points  239  on either side of the images  218 . The process of driving the scatter map from the image is indicated by process block  235  of FIG.  24 . 
     Next as indicated by process block  237 , the scatter map  234  is fit to the normalizing points  222  previously determined at process block  224 . 
     Referring to FIG. 24, the scatter map  234  is thus normalized so that the portions  238  of the scatter map  236  located near the places where the images  218  would fall are given values  222  as determined at process block  224 . This involves a simple shifting up or down of the scatter map  236  and may employ a “least square” fit to shift the scatter map  236  to multiple values  222  obtained from each pin  216 . As adjusted, the scatter map  236  is then subtracted from the image  208  to eliminate or reduce the scatter in that image as indicated by process block  239 . 
     The effect of subtracting a low pass filtered or blurred image properly normalized to actual scatter is to sharpen up the image  208  but also to preserve its quantitative accuracy. Thus the present invention differs from prior art scatter reduction techniques in that it both addresses the variation in scatter across the image caused by attenuation of x-rays by the imaged object but also incorporates accurate measurements of scatter in certain portions of the image. 
     DENSITOMETER ADAPTER 
     Referring now to FIG. 25, a mobile fluoroscopy machine  310  suitable for use with the present invention is similar to that which has been described above with respect to FIG. 1 with exceptions that will be apparent from context. 
     The mobile fluoroscopy machine  310  includes a mobile cart  312  supporting a computer  314  and monitor and keyboard  317  for receiving and processing digital x-ray image data. The cart  312  supports on one side an articulating arm assembly  316  terminating in a rotatable C-arm  318 . The C-arm supports, at the ends of the C, an image intensifier  320  and an x-ray source  322  opposed along an axis  324  so that the x-ray source  322  projects a cone-beam of x-ray radiation toward the image intensifier  320  along axis  324 . 
     The articulating arm assembly  316  is connected to the C-arm  318  through one or more pivotal links  327  so that the axis  324  may be positioned to be horizontal approximately two feet above the floor to rest upon or be supported against the upper end of a supporting pedestal  326  or may be attached to the cart  312 . Referring also to FIG. 26, the pedestal  326  includes a hemicylindrically concave cradle  328  at its upper surface to receive a lower portion of the cylindrical image intensifier  320  when the C-arm is so positioned to rest against the pedestal  326 . 
     The pedestal  326  also provides on its upper surface a channel  330  extending across the axis  324  between the image intensifier  320  and the x-ray source  322  when the latter are positioned on the pedestal  326 . The channel  330  may receive a limb positioner  332  such as may be adapted to support a patient&#39;s foot or arm across the axis  324  for densiometric measurement. The pedestal  326  may be weighted so as to provide a stable surface for support of the x-ray source  322  and image intensifier  320  and to provide adequate support for the patient&#39;s limb. The height of the pedestal  326  is selected to be suitable for either arm or foot imaging. 
     Referring now to FIG. 27 and 30, the channel  330 , extending substantially perpendicularly to axis  324  and has a horizontal bottom surface  333  pierced by two vertically extending guide holes  334  which may be used to receive and position corresponding pins  337  on one of two limb positioners  332 . A foot positioner  336 , as shown in FIG. 26, provides a padded calf support plate  338  fitting adjacent to the bottom surface  333  and an upwardly extending sole support  340  forming an obtuse angle with respect to the calf support plate  338 . A cushion  342  on the calf support plate  338  may be adjusted so as to allow the patient&#39;s leg to extend upward somewhat from vertical for comfort. Gussets  344  span the angle between the sole plate  340  and calf support plate  338  to fix them in relative position but include apertures  346  to allow for the free passage of x-rays through a portion along axis  324  where the os calcis of the heal will be located. 
     When positioned within the channel  330 , the foot positioner  336  is also supported by upwardly extending channel sidewalls  348  which serve further to provide an alignment surface for the imaging face of the image intensifier  320  or other detector array and on the other side, an alignment surface for an emitting face of the x-ray source  322 . Channel sidewalls  348  are generally radio translucent so as to permit the passage of x-rays therethrough, but may include: calibration materials such as are well known in the art for calibrating dual energy devices, antiscatter grids also well known in the art, or occluders for evaluating scatter as have been described above or in the parent applications hereby incorporated by reference. 
     Referring to FIG. 27 and 28, for forearm imaging, the foot positioner  336  is removed and a palm support  352  is inserted by means of pin  337  in one of the holes  334  so as to locate a user&#39;s arm resting against the bottom surface  333  with the user&#39;s palm against the palm support  352  such that the bones of the forearm are placed along the axis  324  for imaging. 
     Referring to FIG. 28 and 29, the hemispherical support cradle  328  may include three radially inwardly extending ribs  354  attached by means of screws or the like to be replaceable. Two of the ribs  354  are positioned in a horizontal plane to substantially bisect the image intensifier  320  when it is placed within the cradle  328 . The third rib  354  is positioned at the bottom of the cradle  328  and is opposed by a rotating locking collar  358  which may be used to further secure the image intensifier  320  within the cradle  328 . The front edge of the image intensifier is abutted against the upright face of the dividing barrier so as to precisely locate it along axis  324 . The inner edges of these ribs  354  define an inner radius  356  of lesser diameter than the cradle  328  that by proper design of the ribs  354  may be adjusted to conform to the outer surface of a particular image intensifier  320 . 
     Referring now again to FIG. 30, some fluoroscopy equipment will not permit digital imaging or the necessary dual energy control needed for densitometry. Accordingly, an independent detector array  360  may be placed within the cradle  328  in lieu of the image intensifier  320 . This detector array  360  may be a pair of stimulable phosphor plates as are understood in the art with intermediate filtering so as to provide dual energy readings with a polychromatic x-ray source. In this way a switching of voltage on the x-ray source  322 , as described above, can be avoided. Alternatively, the detector array  360  may be a large area solid state detector or scanning detector assembly such as are understood in the art including those constructed of amorphous silicon and thin film transistor technology or those employing active pixel technology in which C-MOS integrated circuit fabrication techniques are employed. These detectors may be used with a switched x-ray source  322  to provide dual energy imaging or may be used in a stacked configuration with intermediate filtering so as to provide separate energy measurements, or may be used in a side-by-side configuration with interleaved detector elements filtered so as to be selectively sensitive to different energies. 
     In the preferred embodiment, and as shown in FIG. 31, the independent detector array  360  is a “pancake” image intensifier  361 , suitably small so as to fit within the space between a conventional image intensifier  320  and the x-ray source  322 . Referring to both FIG. 30 and 31, the pancake image intensifier  361  includes a vacuum bowl  362  having a planar front surface  364  for receiving x-rays  366  (normally through the channel sidewalls  348  of the stand  326 ). 
     According to conventional design, the x-rays  366  pass through the front surface  364  of the vacuum bowl  362  to strike a target material  368  to eject electrons  370  into the volume of the bowl  362 . Focusing electrodes  372  direct the electrons to a phosphor  374  where an image is formed to be received by imaging array  375  such as a CCD array or camera. The image area of the phosphor  374  is much smaller than the front surface  364  so as to reduce the image size to one compatible with the camera. In the present invention the distance B between the target  368  and the imaging array  375  (including any optical path through one or more focusing lenses) is less than or equal to the radial dimension. A of the front surface  364  gives the pancake image intensifier  361  an extremely short form factor suitable for practice with the present invention. 
     Hitherto, such form factors were avoided because they are known to result in severe distortion of the image formed on the phosphor  374 . This distortion is accommodated in the present invention by means of digital image processing in computer  314  which receives digitized pixel data from scanning electronics  378  connected to the imaging array  375  and corrects it according to the correction process described above with respect to the pin cushion correction. Accordingly, the addition of digital signal processing allows for production of pancake image intensifier  361  in which the separation of the imaging optics from the front of the image intensifier is much reduced. 
     The above adapter may be modified to use in femur imaging. In this case the  30  pedestal  326  may be eliminated in favor of a positioner (not shown) attached to the image intensifier  320  or x-ray source  322  directly. In the former case, the positioner may provide for a fixed air gap between the patient and the image intensifier  320  to reduce received scatter. So as to allow free manipulation of the C-arm  318 , the positioner may be a lightweight plastic radiolucent material and may optionally include a calibration system such as a flip in phantom for calibration of the dual energy readings and occluders for scatter correction as has been described above. Collimation and/or a separate solid state dual energy image detectors may also be held by the positioner whose outer surface may guide the positioning of the C arm  318  to the necessary orientation which need not be horizontal but may be vertical for fore arm measurements or the like. For femur measurements, the patient may stand and the C-arm  318  manipulated appropriately as guided by the positioner. 
     Referring now to FIG. 31 and 32, computer  314  includes a processor  380  and memory  382 , the latter of which receives raw image data in the form of pixels having spatial locations and brightness values forming images  384 . Memory  382  also includes a processing program  386  providing a general interface and control of the operation of the fluoroscopy machine  310  and a processing of images  384  so as to provide a quantitative measure of bone isolated from soft tissue. 
     The processing program  386  can be simply loaded into the computer  314  for the fluoroscopy machine  310  when the pedestal  326  is to be employed with a fluoroscopy machine  310  providing digital imaging and x-ray voltage control. If an independent detector array  360  is required, the program  386  may be executed on a computer  314  associated with that independent detector array  360 . 
     At a first step in the program  386 , indicated by process block  388 , the operator of the fluoroscopy machine  310 , having indicated a desire to perform densitometry and having positioned the C-arm in the pedestal  326 , enters patient data that will be used to identify the image  384  to be collected. 
     At succeeding process block  390 , data is collected for three distinct images  384  with: 1) no x-ray exposure, 2) high energy x-ray exposure, and 3) low energy x-ray exposure. Each of the exposures is preserved as a separate image file in the memory of the computer  314 . The first exposure is used for correction routines to be described; the latter two exposures are used to deduce bone density according to methods well known in the art in which variations in high energy and low energy absorption are used to deduce the Compton scattering and atomic number of the material lying between the x-ray source  322  and the image intensifier  320 . As is understood in the art, these two measurements allow the amount of bone as opposed to soft tissue located in that image region to be accurately measured. The data is acquired directly from the independent detector array  360  or in the event that stimulable plates are used, a reader may be attached to the computer  314  so as to acquire the necessary pixel data of an image  384 . In the same way a conventional photographic film/filter plate arrangement may be used. 
     At next process block  393 , each of these images is corrected for non-linearity of the detector such as may be determined empirically at an earlier time by testing the detector according to methods well known in the art, and which is a function of the detector and the technology used by the detector. Generally the testing exposes the detector to different fluences of x-rays and measures the output of the detector and the correction is intended to ensure that, for example, a doubling of fluence results in a doubling of detector output after correction. The correction is generally simply a scaling of each of the images by a factor that is a function of the pixel value for each pixel and possibly the location of the pixel. 
     At process block  395 , noise related to the particular line of the detector is removed. Referring to FIG. 34, the imaging array  375  provides a matrix of detector elements  392  arranged in rows and columns. Normally, either rows or columns are ganged together to be read out by dedicated read out electronics  391  spanning a particular row or column. The read out electronics introduces noise which is imposed upon each detector element  392  of that row and which is thus line correlated, that is, more highly correlated with other detector elements  392  of the line than detector elements  392  of different lines. To eliminate line correlated noise, one detector element  392  in each line is blocked by a lead mask  394  so as to be shielded from x-rays. A pixel value  396  from this blocked detector element  392  will provide a value that varies according to the line noise thus a line correlated noise value  398  may be deduced and subtracted from the pixel values  400  of the other detector elements  392  in the line. 
     Referring again to FIG. 32, at a succeeding process block  402 , veiling glare is removed and the field is flattened. This former correction attempts to eliminate blurring of the image such as may be caused by scatter or similar effects within the imaging array  375 . Glare refers generally to a reading that would be obtained under detector elements  392  that were wholly shadowed by an occluding absorber on the surface of the array  375 . The glare is a function of the detector technology and is reduced by a deconvolution process based on an empirically derived deconvolution kernel according to a number of techniques well known in the art. 
     Also at process block  402  the field is flattened which is to say the gain variation of the detector elements  392  are normalized according to an empirically derived normalization map determined at the factory by exposing the detector to a uniform x-ray elimination and noting variation and intensities reported in the pixel values  400 . At this time, dark currents from the detector elements  392  may also be eliminated as determined from the no-exposure x-ray image taken at process block  390 . 
     Referring now to process block  404 , a dynamic scatter correction may be employed as has been previously described with respect to FIGS. 21-23. Alternatively referring also to FIG. 33, a dynamic scatter correction may be employed in which the data of the image  384  is analyzed so as to create a histogram  406  of pixel values  400  for the entire image. The histogram may be divided into regions  408  (five equal regions in the preferred embodiment) corresponding roughly x-ray paths through: 1) air-only, 2) thin tissue, 3) thick tissue only, 4) thin bone and, and 5) thick-bone. Each of these materials will exhibit a different scattering and hence a different empirically derived scatter kernel  410  may be assigned to each region  408  with generally the lower density regions having narrower kernels commensurate with less scatter. 
     The selected scatter kernel  410  may be scaled by the pixel value of the image  384  on a pixel by pixel basis and that kernel, so scaled, applied to a deconvolver  412  used to deconvolve the image  384  to produce a deconvolved image  414 . A number of techniques of deconvolution are well known in the art using a fixed scatter kernel and these same techniques may be used with the variable scatter kernel  410  described here. During deconvolution the kernel  410  will be sequentially applied to a set of adjacent pixel values determined by the width of the kernel. The center pixel value at any step of the deconvolution will be used to scale the kernel and to identify the region of the histogram for the purpose of selecting the kernel  410 . In an alternative embodiment, the kernel may be fixed and simply scaled by the value of the centermost pixel during de-convolution. 
     Referring again to FIG. 32, at process block  416 , the images  384  are log corrected reflecting the fact that attenuation is exponentially related to thickness. The images are now related to thickness, a dimension which will be important in the ultimate bone-density determination. 
     At following process block  418 , speckle may be identified for certain x-ray detectors  360  that are subject to extremely high readout values caused by noise which is possibly related to direct x-ray irradiation of the detector element. Speckle is identified by a simple thresholding process. 
     At next process block  420 , path length correction may be performed based on the geometry of the particular C-arm such as may vary path length and magnification across the image as is well understood in the art. 
     Similarly at succeeding process block  422 , beam hardening, the well known effect of a spectral shift in a polyenergetic x-ray beam as it passes through different thicknesses of material, and a Heel effect correction may be made, the Heel effect correction referring to a variation in the spectrum of an x-ray beam as a function of its angle in the cone of x-ray beams. Both of the corrections are known in the art, but must be employed in the present invention in order to provide suitable quantitative accuracy for densitometry. 
     At process block  424 , the identified speckle of process block  418  is corrected by eliminating these identified pixels from subsequent calculation or by replacing them with a local average value. 
     The entire image may then be averaged or low-passed filtered at process block  426  so as to further reduce noise and to eliminate unneeded resolution. 
     The images are then processed according to well understood techniques to produce a bone mineral density value at process block  428 . This bone mineral density value indicates the amount of bone material at each pixel of the image largely independent of surrounding soft tissue. The pixel image may be analyzed in a number of methods but most simply, as indicated by process block  430 , by defining either automatically or manually a desired region of interest within the image and making a measurement of total bone density within that region. Automated techniques may look for a local maximum or minimum of bone density or may use image recognition type techniques to locate reproducibly a particular region of the forearm or os calcis. Morphometric analysis may be applied to the image to detect bone fracture and other techniques such as texture analysis may be performed according to methods well known in the art. The results of the analyses and images so processed may be displayed by the computer  314 . 
     It is thus envisioned that the present invention is subject to many modifications which will become apparent to those of ordinary skill in the art. Accordingly, it is intended that the present invention not be limited to the particular embodiment illustrated herein, but embraces all such modified forms thereof as come within the scope of the following claim.