APPARATUS AND METHOD FOR REAL-TIME TRACKING OF TISSUE STRUCTURES

Abstract

A method and system are disclosed for radiosurgical treatment of moving tissues of the heart, including acquiring at least one volume of the tissue and acquiring at least one ultrasound data set, image or volume of the tissue using an ultrasound transducer disposed at a position. A similarity measure is computed between the ultrasound image or volume and the acquired volume or a simulated ultrasound data set, image or volume. A robot is configured in response to the similarity measure and the position of the transducer, and a radiation beam is fired from the configured robot.

Claims

1. A system for estimating ultrasound image quality at a variety of positions to determine a desired ultrasound transducer position, the system comprising: an input module configured to receive anatomical image data from a first imaging modality, the anatomical image data including a target; an image quality estimation module coupled to the input module, the image quality estimation module configured to estimate ultrasound image quality when imaging the target at a position relative to the target by analyzing the acquired image data to estimate ultrasound velocity along a route from the position to the target, the route passing through one or more tissue types; and a display module coupled with the image quality estimation module, the display module configured to output an indication of the ultrasound image quality at positions relative to the target to the user.

2. The system of claim 1, wherein the received image data comprises CT image data, and wherein the route is evaluated by the image quality estimation module by classifying the route tissue types based on their CT intensity values.

3. The system of claim 2, wherein the image quality estimation module virtually propagates an ultrasound beam along the route and through the one or more tissue types using a ray casting method and estimates ultrasound transmission at discrete sampling points by calculating a different between incoming beam strength and tissue absorption and adjusting for reflection.

4. The system of claim 1, wherein the image quality estimation module is configured to estimate ultrasound image quality over a time period.

5. The system of claim 4, wherein the input module is further configured to receive ultrasound image data, and the system further comprises a signal processing module coupled with the input module, the signal processing module configured to compensate for distortion in the ultrasound images by a constant or dynamic gain factor and a constant or dynamic offset factor.

6. The system of claim 5, wherein the signal processing module is configured to use the constant gain factor and the constant offset factor when the estimated ultrasound velocity along the route from the position to the target is assumed to not change over the time period.

7. The system of claim 5, wherein the signal processing module is configured to use the dynamic gain factor and the dynamic offset factor when the estimated ultrasound velocity along the route from the position to the target varies over the time period.

8. The system of claim 1, wherein the output of the display module comprises a three dimensional model of a surface of the anatomical site with color variations.

9. The system of claim 1, wherein: the received image data comprises CT image data; the route is evaluated by the image quality estimation module by classifying the route tissue types based on their CT intensity values; and the image quality estimation module virtually propagates an ultrasound beam along the route and through the one or more tissue types using a ray casting method and estimates ultrasound transmission at discrete sampling points by calculating a difference between incoming beam strength and tissue absorption and adjusting for reflection.

10. An ultrasound tissue imaging system comprising: an ultrasound image sensor to generate ultrasound data of a target included within or adjacent to an anatomical site; an input for receiving other image data regarding the anatomical site from another imaging modality; a processor configured to generate ultrasound image correction data of the anatomical site from the other image data to compensate for distortion of the ultrasound data; and an image or measurement output coupled to the processor for communication of an ultrasound measurement or image of the target in response to the correction data.

11. The system of claim 10, wherein the other image data comprises CT image data.

12. The system of claim 10, wherein the another imaging modality comprises a bi-plane x-ray imaging modality.

13. The system of claim 10, wherein the anatomical site is a heart.

14. The system of claim 10, wherein the ultrasound image correction data is generated less frequently than the ultrasound data.

15. The system of claim 10, wherein the processor generates the ultrasound image correction data by observing a time sequence of positions of the target using the another imaging modality.

16. The system of claim 15, wherein the processor develops a first motion curve of the target from the ultrasound data and a second motion curve of the target from the another imaging modality.

17. The system of claim 16, wherein the processor is further configured to deform the first motion curve by overlaying the first motion curve with the second motion curve and compute the distortion in the ultrasound data based on a deformation of the first curve.

18. The system of claim 17, wherein the processor is further configured to subtract the distortion from the ultrasound data.

19. The system of claim 1, wherein the processor is further configured to extrapolate in-treatment position of the target based on past observations.

20. The system of claim 1, wherein the ultrasound measurement or image of the target comprises a three dimensional model of a surface of the anatomical site with color variations.

Description

BRIEF DESCRIPTION OF THE DRAWINGS

[0027] FIG. 1 illustrates an exemplary method of taking ultrasound images according to embodiments of the invention;

[0028] FIG. 2 illustrates an exemplary method of estimating ultrasound image quality according to embodiments of the invention;

[0029] FIG. 3 provides an exemplary table mapping Hounsfield values to acoustic properties;

[0030] FIG. 4 illustrates an exemplary method of estimating ultrasound image quality according to embodiments of the invention;

[0031] FIGS. 5A-5B illustrate a display of the output of ultrasound image quality estimates calculated according to methods of the present invention relative to preferred ultrasound transducer positions identified by an ultrasound technician;

[0032] FIG. 6 illustrates a display of the output of ultrasound image quality estimates calculated according to methods of the present invention for imaging the substantia nigra;

[0033] FIGS. 7A-7E illustrate optimal windows for visualizing the RUPV for different patients according to methods of the present invention;

[0034] FIG. 8 illustrates optimal windows for visualizing the septum for eight time steps according to methods of the present invention;

[0035] FIG. 9 illustrates a situation where estimated ultrasound velocity along a route from the ultrasound transducer position to the target may vary over time;

[0036] FIG. 10 illustrates a situation where estimated ultrasound velocity along a route from the ultrasound transducer position to the target stays relatively constant over time;

[0037] FIG. 11 illustrates an exemplary method of compensating for distortion in ultrasound image data;

[0038] FIG. 12 illustrates a first motion curve fitted to ultrasound data relative to a second motion curve fitted to intermittently acquired image data from another imaging modality;

[0039] FIG. 13 illustrates an exemplary method of extrapolating target position based on past tissue motion; and

[0040] FIG. 14 illustrates a system for estimating ultrasound image quality at various positions and/or compensating for distortions in ultrasound image data.

DETAILED DESCRIPTION OF THE INVENTION

[0041] The present invention generally provides improved devices, systems, and methods for ultrasound tracking of tissue. The invention is particularly well suited for tracking of moving tissues such as tissues of the heart and tissue structures adjacent the heart that move with the cardiac or heartbeat cycles. The invention may take advantage of radiosurgical structures and methods which have been developed for treating tumors, particularly those which are associated with treatments of tissue structures that move with the respiration cycle. The systems and methods disclosed herein may be used to continuously track movement of a patient's heart during radiosurgical examinations for example. The cardiac cycle is typically considerably faster than the respiration cycle. The overall treatment times can also be quite lengthy for effective radiosurgical procedures on the heart (typically being greater than 10 minutes, often being greater than ½ hour, and in many cases, being two hours or more). Hence, it will often be advantageous to avoid continuous imaging of the target and adjacent tissues using fluoroscopy or the like. A variety of differing embodiments may be employed, with the following description presenting exemplary embodiments that do not necessarily limit the scope of the invention.

[0042] FIG. 1 illustrates an exemplary method 10 of estimating ultrasound image quality at ultrasound positions relative to a target. At step 12, pre-treatment images of the target are acquired, typically over a time period. At step 14, the pre-treatment images may be used to estimate image quality of an ultrasound transducer at various positions relative to a target. The image quality may be estimated over the time period. At step 16, the image quality estimates for one or more positions may be displayed to an operator, preferably relative to a three-dimensional surface model of the patient. At step 18, a desired transducer position may be identified which has a high estimated image quality. The desired transducer position may have the highest averaged image quality over a given time period. At step 20, ultrasound velocity along the route from the desired transducer position to the target may be analyzed to determine whether the ultrasound velocity along the route stays relatively constant or whether it varies with time. At step 22, ultrasound images of the target may be acquired from the desired position. At step 24, distortion in the ultrasound images may be compensated for using a gain and offset factor calculated from the information obtained in step 20. Advantageously, an ultrasound operator may be able to quickly identify and visualize the optimal positions for an ultrasound transducer prior to imaging a target by utilizing the disclosed method. Further, the ultrasound operator may be able to visualize how the image quality may change from a plurality of positions with patient movement from a respiratory motion and/or a heartbeat motion. Moreover, distortions in the acquired ultrasound images may be compensated by using the methods disclosed herein.

[0043] The pre-treatment image data acquired 12 include the anatomical site and the target tissue. The image data may be a two dimensional single CT scan or a three dimensional CT scan. Further the pre-treatment image data may be acquired over a time period in order to capture anatomical site and target movement during a heartbeat cycle and/or a respiratory cycle motion of the patient. These scans may also be used for the contouring of the target region and the calculation of the dose distribution during radiosurgical treatment planning. The scans may come from other imaging modalities such as MRI, PET, etc.

[0044] FIG. 2 illustrates an exemplary method 26 of estimating the image quality of a transducer placed at a certain skin position 14. The method starts at step 28 by identifying an ultrasound transducer position relative to the target in the acquired image data. At step 30, a route from the transducer position to a chosen target position inside the volume is analyzed. At step 32, the ultrasound velocity in tissues along the route is estimated. At step 34, an ultrasound beam may be virtually propagated from the position and along the route based on a ray casting method. At step 36, ultrasound transmission may be estimated at discrete sampling points to estimate overall ultrasound image quality.

[0045] To estimate the imaging quality of a transducer placed at a certain position, the route from the transducer position to a chosen target position inside the volume must be analyzed. For a good view of the target one consideration is the absorption of the ultrasound beam travelling forth and back between the target and the transducer. The estimation of beam strength can be calculated based on beam transmission and reflection as the main parts in beam absorption. FIG. 3 illustrates an exemplary pre-treatment scan 38 of the anatomical site including the target 40. A route 42 from a transducer position 44 to the target 40 is determined. The tissues which the route 42 passes through can be determined from the pre-treatment image data. Tissue types may be classified by using the acquired pre-treatment image data. For example, where the pre-treatment data is one or more CT scans, the tissue types may be classified with respect to their CT intensity values. FIG. 4 provides an overview of tissue in the region of the heart and their corresponding Hounsfield units. The exemplary route 42 illustrated in FIG. 3, passes through skin 46, fat 48, and muscle 50 in order to arrive at target 40 from the transducer position 44. Afterwards, the ultrasound beam is virtually propagated through the CT volume using a ray casting method. For every discrete sampling point 52, the ultrasound transmission is calculated as a difference of incoming beam strength and tissue absorption. The reflection may be calculated by:

[00001]$\begin{array}{cc}R={\left(Z_2-Z_1\right)}^2/{\left(Z_2+Z_1\right)}^2& \left(1\right)\end{array}$

Where Z.sub.1 and Z.sub.2 represent the acoustic impedances at the current and previous sampling point, is additionally subtracted. Given all possible transducer positions 44 on the patient's chest, a desired ultrasound transducer position maximizes the ultrasound transmission between transducer and target. In some embodiments, the target visibility over time may be calculated where the desired position is computed as the weighted maximum visibility over all time steps of a 4DCT. Additionally, embodiments may take advantage of enhanced beam strength computation by adjusting for interference, refraction, diffraction and beam diffusion. Further, methods may include additional visibility analysis by analyzing target reflection, entropy in target region, and complete ultrasound image simulation. The methods may take a simulated ultrasound image and detect, at which quality the target is visualized inside. The easiest way to do it may be to accumulate the brightness in this area. An enhanced method may be to measure the entropy or the amount of edges/intensity gradients inside the image. All gradients could be summed to one value indicating the image quality.

[0046] After image quality estimation 14, the results may be displayed to an operator 16. Preferably, a plurality of image quality estimations 14 are calculated according to the methods disclosed herein. Indications of the image quality at each position may inform an operator which positions are best for imaging a target using an ultrasound transducer. In a preferred embodiment, the indications of image quality are displayed relative to a three dimensional model of the patient. The image quality indications may comprise a color coding on the three dimensional model. For example, FIG. 5A displays the computed image quality of transducer positions according to methods of the present invention. As shown in FIG. 5A, a three-dimensional model 54 of the patient can be constructed from the pre-treatment image data. The blue indications 56 on the three-dimensional model indicate positions which have high estimated image quality. FIG. 5B shows the ultrasound windows as determined by a radiologist. The positive windows 58 illustrate the preferred windows for measuring the target tissue while the negative windows 60 illustrate undesired windows. As can be appreciated from FIGS. 5A and 5B, the methods of estimating ultrasound image quality prior to imaging disclosed herein correspond to the positive windows 58 which were determined by the radiologist. It may be worth noting that positive window 62 which was added by the radiologist required a lot of pressure to obtain a reasonable image and accordingly, the ultrasound image quality estimation did not include a corresponding positive window.

[0047] FIG. 6 illustrates another exemplary display of image quality estimates as determined by methods disclosed herein. The high quality windows for imaging the substantia nigra is displayed. This type of imaging may provide for the early detection of Parkinson's disease. FIGS. 7A-7E illustrate the optimal viewing windows for visualizing the RUPV for different patients as computed by the methods disclosed herein.

[0048] As discussed above, image quality estimation may be performed for a plurality of time steps to indicate how image quality may change with a patient respiratory and/or heart beat motion. FIG. 8 illustrates the optimal windows for visualizing the septum over eight time steps. The optimal windows may vary slightly over the eight time steps. Image quality may also be estimated at varying frequencies as needed. For example, the quality may be estimated at higher frequencies for target tissues which have fast tissue movements such as the heart. Image quality may be estimated at lower frequencies for tissues which relatively slower motions. Accordingly, the operator may choose a transducer position with the highest average image quality over the time period for visualizing a target tissue.

[0049] After estimating the ultrasound image quality at one or more positions 14, steps maybe taken to minimalize distortion or to compensate for distortion in any acquired ultrasound images. The distance in ultrasound scans may be computed as:

[00002]$\begin{array}{cc}\mathrm{Distance}=\mathrm{Time}*\mathrm{Velocity}& \left(2\right)\end{array}$

This ultrasound velocity cannot be measured by the ultrasound transducer and is commonly approximated by 1540 m is as a mean speed of sound for the penetrated tissue. While this maybe a good approximation for visualization, it theoretically allows for errors of up to five percent in the distance calculation. In a worst-case scenario, a target at a distance of 150 mm may be distorted in forth/back direction by up to 7.5 mmm including an uncertainty of 15 mm. While the theoretical error is relatively unlikely, distortions of up to two percent (e.g., 3 mm at 150 mm distance) are commonly present.

[0050] In some embodiments, distortions can be derived from the pre-treatment image data, such as a CT volume, and used for compensation. Preferably, an optimization of the transducer position may be used as an a priori strategy to minimize the expected velocity changes and the necessity to dynamically correct for them.

[0051] As seen above, tissue ultrasound properties can be estimated based on acquired pre-treatment data. For example, when pre-treatment data comprises CT image data, ultrasound properties can be sufficiently mapped to CT Hounsfield units for a particular anatomical region. An ultrasound transducer may be virtually placed at the skin surface and the acoustic properties including velocity differences can be computed from CT data. Assuming constant velocity errors between transducer and target, a simple function using the two compensation factors, Gain and Offset can be applied to correct for distortions:

[00003]$\begin{array}{cc}\mathrm{Corrected}\mathrm{distance}=f\left(\mathrm{Distance},\mathrm{Gain},\mathrm{Offset}\right)& \left(3\right)\end{array}$

[0052] The distortion of the relative target movement may be low and may be neglected. The gain and offset may be calculated with the provided set of two corresponding positions/distances in ultrasound and a secondary modality. Then a function may be used to linearly undistort the ultrasound volume:

real_distance=f(ultrasound_distance)=offset+gain*ultrasound_distance.  (4)

[0053] In situations with time-varying errors due to non-uniform movements in the beam path, the absolute and even the relative target movements can be compensated. FIG. 9 illustrates a situation where a movement along the beam path induces a non-existing target movement in the distance measurements. Ultrasound transducer 64 may be imaging target 66 from a route 68. Route 68 passes through a tissue 70 which moves during a patient's respiration and/or heartbeat motion. Accordingly, even though target 66 is static, the measured target distance 72 may vary over time due to the motion of tissue 70.

[0054] The quality of the compensation depends on the available data. In the case of only one CT volume, static distortions may be compensated for quite well. Error sources may include errors from mapping ultrasound properties according to CT Hounsfield units and ultrasound model complexity. Compensation for static error can include a measurement of static errors using the CT image data and a static offset and gain compensation. Resulting image quality may be very good. In situations where the target is moving, it is preferable to have a series of pre-treatment scans available to compensate for errors. It may be difficult to calculate for distortion since target motion is unknown with only a single CT volume. In situations where only a single CT volume is available, a statistical approach may be used to compensate for distortion. Alternatively, the assumption of the common mean ultrasound velocity may be used. When 4DCT data is available, error sources may include errors from mapping ultrasound properties according to CT Hounsfield units and ultrasound model complexity. A further error source may be introduced in the identification of the current time step. Error may be compensated by a measurement of static errors over all time steps. Thereafter, a dynamic gain and offset compensation may be calculated to reduce distortions due to respiratory and/or cardiac motion. Accordingly, ultrasound image data distortion may be compensated when the target is static or when the target is in motion by using the acquired pre-treatment data.

[0055] Another method of minimizing distortion in ultrasound transducer image data comprises minimizing the distance between the ultrasound transducer and the target. The distance may be simply added as a factor into the quality function of the search algorithm given above.

[00004]$\begin{array}{cc}\mathrm{Position}\mathrm{quality}=a*\mathrm{Transmission}+b*\mathrm{Distance}& \left(5\right)\end{array}$

[0056] When 4DCT data is available, minimizing distortion may include measuring the absolute distortion or the distortion change (combined with static distortion correction) overtime for every possible transducer position and minimize them in order to find an optimal transducer position.

[00005]$\begin{array}{cc}\mathrm{Position}\mathrm{quality}=a*\mathrm{transmission}+b*\mathrm{Distortion}\mathrm{Change}& \left(6\right)\end{array}$

[0057] When CT image data of a single time step is available, transducer positions where tissue velocities are similar for a cylindrical volume around the beam path may be preferred. FIG. 10 illustrates such an example. In this situation, transducer 74 may be imaging target 76 along route 78. A tissue 80 with a different ultrasound property may have a motion associated with patient respiratory or heartbeat motion, however the tissue 80 and its motion may be disposed outside of a security region 82. Accordingly, the transducer position may be selected to make it less likely that tissue with different velocity properties enters the beam path and induces time-varying distortions. Thus the methods disclosed herein may be used to select ultrasound transducer positions which limit the amount of distortion in image data by limiting the chances of a time-varying ultrasound velocity along the ultrasound transducer route.

[0058] As discussed above, the methods disclosed above may be used to track target tissues during radiosurgical treatments and systems such as those disclosed in related U.S. patent application Ser. No. 11/971,399 entitled Depositing Radiation in Heart Muscle under Ultrasound Guidance, the entire disclosure of which is incorporated herein by reference. A stereotaxic radiation treatment device typically includes a beaming device, which is also called the beam source. This beam source produces a beam of radiation. The beam source can be moved to different locations. In this way, the beam can be directed towards the target. Targets are typically tumors, but other lesions can also be treated with this method. In commercially available systems for radiosurgery, the beam source is mounted to a robotic arm. This arm is freely programmable, and can move the beam source to appropriate locations in space. It can also move the beam source in such a way that the beam tracks the motion of a moving target. The motion of the target occurs when the tumor is close to the heart or the lung, and is due to the heartbeat or the respiratory motion of the patient's chest. Prior to treatment, a CT or an MR may be taken from the anatomical site/region of interest. The target is then marked in the resulting stack of images or may be marked relative to a three-dimensional image model of the target region according to the methods and systems disclosed in related application Ser. No. 12/838,308 entitled Heart Tissue Surface Contour-Based Radiosurgical Treatment Planning, the entire disclosure of which is incorporated herein by reference.

[0059] Radiosurgical systems typically track target tissue using a bi-planar x-ray system, however there are advantages of reducing the amount of patient exposure to radiation. While ultrasound tracking of target tissues has been proposed, distortion in ultrasound image data may make such tracking insufficient since target tracking during radiosurgical treatments must be very accurate, especially for target tissues, such as heart tissues, which move relatively rapidly with patient respiration motion and/or heartbeat motion.

[0060] FIG. 11 illustrates another method 84 of reducing distortions in acquired ultrasound imaging data and/or for tracking target tissues during a radiosurgical procedure. At step 86, markers may be implanted at or adjacent to the target tissue if the target tissue cannot be easily distinguished in imaging data. At step 88, the position of the target tissue is determined using a first imaging modality to image the target tissue and/or the implanted markers at discrete intervals. At step 90, the position of the target tissue is determined using an ultrasound imaging device to image the target tissue and/or the implanted markers to provide a continuous stream of imaging data. At step 92, a first curve is fitted to the target tissue position data at the discrete intervals as determined by the first imaging modality. At step 94, a second curve is fitted to the target tissue position data as determined by the ultrasound imaging device. At step 96, the first curve and the second curve are overlaid to obtain a single curve. At step 98, distortion in the ultrasound image data may be subtracted based on the deformation of the ultrasound curve from the overlaying of the two curves. At step 100, current positions of the target tissue may be extrapolated based on past motion during times of poor ultrasound imaging.

[0061] In some methods of the present invention, markers may be implanted near a target 86. These markers may be, for example, gold markers which are implanted in or near a target tissue. The markers may be used to better identify a target tissue in subsequently captured images. The exact position of the target tissue and/or implanted markers may be determined by a first imaging modality 88. The first imaging modality may be a stereo x-ray imaging system which typically accompanies radiosurgical systems. Continuous imaging is discouraged due to radiation exposure of the patient and technical limitations of the x-ray imaging systems available commercially. The acquisition of images with the stereo x-ray imaging system may be repeated several times before treatment. This gives a series of exact positions of the target obtained with x-ray imaging. However, this series may not be continuous and may not present real-time information on the target location. Accordingly, the ultrasound camera may be used to yield a continuous stream of images 90, where each image has the target located. As discussed above, the target location may be identified with high speed automatically by using a template matching algorithm. This stream may contain more than 20 images per second.

[0062] A first curve may be fitted to the position data obtained by the stereo x-ray imaging system 92 and a second curve may be fitted to the position data acquired by the ultrasound image data 94. Since the second curve is obtained with ultrasound, it will be subject to distortion. The first such curve is free of distortion. After fitting the first and second curves to the acquired position data, the two curves are overlaid to obtain a single undistorted curve 96 with points sufficiently densely spaced along the curve such that real-time tracking of the target is possible.

[0063] FIG. 12 illustrates the two curves representing position information for the target. For example, each curve can represent the motion of the target along the x-axis of the base coordinate system used for the treatment. The curve 102 shows the distorted curve of target positions obtained from analyzing the ultrasound images. The curve 104 may not be fully known during treatment, and only a small number of positions 106 along this curve are known. Since both curves must show the same motion, they can be aligned. After alignment, the ultrasound curve 102 is deformed in such a way that the points 106 will closely match the curve 104. The deformation thus obtained is then expressed as a function and can be used to obtain undistorted position information 98 for the new ultrasound images received after switching of the x-ray imaging device. An ultrasound image may be acquired along the beam path from transducer to target and back. Distortions may be mostly caused by differences of the speed of sound in different tissue. This may be a non-linear distortion along the beam path and (to have the best fitting) a non-linear function:

f(ultrasound_distance)=real_distance  (7)

has to be found for undistortion. This may be a table, which can be approximated by a polynomial function.

[0064] An ultrasound imaging camera must be placed in a position with good visibility on the patient's skin surface. In a preferred embodiment of the present invention, the method includes obtaining an optimal placement of the ultrasound camera 10. As discussed above, a simulation of the physical properties of the ultrasound imaging device is used. This simulation may rely on the CT or MR data obtained before treatment.

[0065] In another preferred embodiment in accordance with the present invention, the past curve of the motion obtained with the ultrasound system is used to extrapolate current and future positions 100. This may be done to overcome periods of poor visibility in the ultrasound image. FIG. 13 shows several positions of a target as seen in an ultrasound image. In each image, the target occurs shifted, either in vertical or horizontal direction. A template may be delineated by the user in the first image 108. Cross correlation is used to compute the shifting offsets 110 and 112 for obtaining the new position of the target.

[0066] In a further embodiment of the invention, the x-ray imaging system is used to locate small markers attached to the ultrasound camera. In this way, the relative positions of the ultrasound camera can be computed in the same coordinate system as that of the treatment device.

[0067] FIG. 14 illustrates a system 114 for compensating for distortions in ultrasound images. The system 114 may include an input module 116 for receiving: 1) pre-treatment images of an anatomical site including the target from a pre-treatment imaging modality 118; 2) ultrasound images of the anatomical site including the target from an ultrasound imaging modality 120; and 3) in-treatment images of the anatomical site including the target from an in-treatment imaging modality 122. An image quality estimation module 124 may be coupled with the input module 116 for estimating ultrasound image quality at various positions relative to the target. A signal processing module 126 may be coupled to the input module 116 for adjusting for distortions in acquired ultrasound image data by using acquired pre-treatment image data or in-treatment image data. The output from the image quality estimation module 124 and the signal processing module 126 may be displayed to an operator by using a coupled display module 128. In a preferred embodiment, system 114 includes an extrapolation module 130 for calculated current or future target position based on past target motion.

[0068] The image quality estimation module 124 may use acquired pretreatment images of the anatomical site to estimate ultrasound image quality at a plurality of positions to determine a desired ultrasound transducer. The image quality estimation 124 module may estimate image quality over a time period, such as a time period over a patient respiratory motion and/or heartbeat motion. The signal processing module 126 may be configured to compensate for distortions in acquired ultrasound image data by using methods disclosed herein and acquired pre-treatment images of the target area or acquired in-treatment image data, such as stereo x-ray images. The display module 128 may be configured to output the ultrasound image quality estimates relative to a three dimensional model of the patient.

[0069] While the exemplary embodiments have been described in some detail, by way of example and for clarity of understanding, those of skill in the art will recognize that a variety of modification, adaptations, and changes may be employed. Hence, the scope of the present invention should be limited solely by the appending claims.

[0070] The embodiments discussed herein are illustrative. As these embodiments are described with reference to illustrations, various modifications or adaptations of the methods and/or specific structures described may become apparent to those skilled in the art after reading the above disclosure.

[0071] In the foregoing specification, the invention is described with reference to specific embodiments thereof, but those skilled in the art will recognize that the invention is not limited thereto. Various features and aspects of the above-described invention can be used individually or jointly. Further, the invention can be utilized in any number of environments and applications beyond those described herein without departing from the broader spirit and scope of the specification. The specification and drawings are, accordingly, to be regarded as illustrative rather than restrictive. It will be recognized that the terms “comprising,” “including,” and “having,” as used herein, are specifically intended to be read as open-ended terms of art.