fig1 is a schematic drawing of an ultrasound probe 10 which may be used to take a transesophageal echocardiogram ( tee ). such probes 10 are adapted to scan a fan - shaped image slice , as depicted in fig2 . when rotating the probe 10 around its axis 12 , as shown by arrow 14 , a number of radial slices 11 may be acquired . the field of view 16 of such a probe 10 is therefore approximately cone - shaped . in the so - called near - field 17 , i . e . the area nearer to the probe 10 , the ultrasound rays converge , and therefore the data points corresponding to neighbouring voxels are closer together . this is exemplified in fig2 by pixel 21 . in the area further away from the probe 10 , the so - called far - field 19 , the ultrasound rays diverge , and therefore the data points corresponding to one voxel are further apart , as exemplified by voxel 22 in fig2 . please note that the 2d image dataset of fig2 may be generated with any kind of ultrasound probe adapted for 2d imaging , not only the tee probe 10 shown in fig1 . thus , the inventive method may be applied to any type of ultrasound image data . fig3 demonstrates further the principle of rotating ultrasound : a cone - shaped 3d dataset 24 can be generated for example with tee probe 10 according to fig1 . for visualization , a plane 26 is placed through the cone - shaped dataset 24 , perpendicular to axis 12 , and the data points along plane 26 are extracted . the image slice 28 hereby generated is shown in fig4 . the 2d image dataset 28 is essentially circular and is build up by image data spaced along radial lines 11 , which ideally should all cross the centre point p . as will be explained in more detail below , it can be advantageous to apply the method of the invention on this type of dataset 28 . according to the inventive method , a filtering algorithm is applied on the input dataset , which may have a format as shown in fig2 or 4 , or any other format . preferably , the filtering algorithm is a kernel smoother , whereby each data point is averaged with its neighbours in the image according to a certain kernel function . fig5 and 6 show examples of suitable kernels . please note that , while fig5 and 6 show 1d kernels , the same kernel functions may be 2d or 3d . for this reason , the figures show the kernel functions depending on an unspecified variable x , which may be any spatial variable such as x , y , z in cartesian coordinates , or r in polar coordinates . usually , the kernel function will be isotropic . furthermore , since the kernel is applied to a dataset with discrete data points , it is preferably also discrete . ideally , the sum of every value of the kernel function will be 1 , since this avoids any normalisation during the filtering step . fig5 shows an example of a 1d gaussian function φ ( x ) with 9 data points . below − 4 and above 4 , the value of the kernel function will be 0 . thus , a filtering or smoothing step using this kernel will result in an averaging over the nearest 9 pixels . fig6 shows an example of a box function b ( x ) with 5 data points . if this function is in 2d , it should be 5 × 5 box function , wherein each point has a value of 0 . 04 , so that the sum of all data points is 1 . however , please note that other kernel functions may also be used . fig7 shows an overview of a method according to preferred embodiment of the invention . an input image dataset 30 is received from an input source , e . g . from an ultrasound machine , from a data storage ( offline ) or from a network . the input image dataset is filtered in step 31 , preferably with a kernel filter with a fixed kernel size , e . g . a gaussian smoothing filter with a kernel size of 5 × 5 in the azimuthal and elevational direction . thereby , a filtered image data set 32 is generated . the original ( unfiltered ) input image dataset is kept . at 34 , a spatially varying weighting function between the filtered and the input image dataset is either computed , or provided from some input source . since this function is simple to evaluate , e . g . a linear , quadratic or other polynomial ramp between the near - field and the far - field , it can be computed voxel by voxel in real time . in step 33 , the input image dataset and the filtered image dataset are added , thereby using the spatially varying weighting function 34 , e . g . using the formula cited above . the result of this weighted addition 33 is an output image dataset 36 having the same data structure as the input image dataset 30 and the filtered image dataset 32 . if a suitable weighting function 34 has been used , the output image dataset 36 may for example achieve good smoothing in the near - field , while keeping the detail in the far - field . optionally , surface or volume rendering is applied on the output image dataset 36 in step 37 . the result of the rendering is visualized e . g . on a screen in step 38 . the whole method as shown in fig7 may be carried out in real time . this means that the practitioner may view the filtered and rendered image immediately after acquiring the image data , without any noticeable waiting times in - between . the filtering method of the invention can be used for black and white image data or for colour image data . two different preferred applications of the invention shall now be described with reference to fig8 - 10 and 11 - 14 , respectively . fig8 shows a 2d fan - shaped input image dataset , which may for example be generated by ultrasound . the slice spans an angle of θ , and the radial dimension ( range ) is shown as r . the raw image data are in acoustic coordinates , which has the consequence that pixels such as pixel 21 close to the origin are small , while pixels 22 in the far - field cover a larger area in real space . thus , the image dataset 20 has a spatially varying pixel size . to demonstrate the consequences of this spatially varying pixel size on the filtering algorithm , the same dataset 20 is shown in fig9 spread onto cartesian coordinates , but without a re - sampling step . therefore , the structure 44 , which is shown as a circle of dots , is severely distorted . if a standard filtering algorithm using a fixed size kernel would be applied to the dataset of fig8 , it would be impossible to select a kernel size which is optimal for both the near - field and the far - field : if the filter is set up such that the near - field , ( e . g . pixel 21 ) looks good , the far - field would be too blurred , due to the large kernel size . on the other hand , if the filter is designed for the far - field to look good , the smoothed image is often not enough smoothed in the near - field . therefore , the invention provides a method wherein different degrees of smoothing may be applied to different areas of the image dataset , while still using a smoothing algorithm which is essentially unvarying over the dataset , e . g . by using a fixed - size kernel . for example , a box - size kernel with size 3 × 3 may be applied on dataset 20 . a suitable weighting function w ( r ) for this dataset is shown in fig1 : the weighting function w ( r ) is independent of angle θ , but has a quadratic fall off from the near - field to the far - field , i . e . it is proportional to 1 / r , wherein r is the range direction . this is because we have quadratic kernel 3 × 3 , the weighting function should also vary quadratically . the first data point w ( r 1 ) of the weighting function corresponds to the first pixel row in dataset 20 , which is at a certain distance r 1 from the origin . fig1 shows a different type of dataset 28 , which is the type of circular image plane which may be obtained by putting a slice through a cone - shaped dataset 24 acquired by means of rotating ultrasound , see fig3 and 4 . as mentioned above , the dataset is acquired slice by slice , while the probe 10 is rotated e . g . by a small electric machine . however , during the acquisition period , the subject will usually move , so that the slices s 1 , s 2 , . . . , s 8 will not be ideally located , and in particular will not all cross the centre point p . furthermore , it is possible that the electric machine is not precise , so that the angles between adjacent slices will not be exactly identical . however , these deviations cannot be traced , so that the observer has to assume that the centre point of each slice s 1 , s 2 , . . . , s 8 will be at the position of point p . to alleviate the artefacts created by these impositions , one should apply a comparatively strong smoothing or filtering algorithm in the centre of dataset 28 . the peripheral regions of dataset 28 , by contrast , can be filtered to a lesser degree . this is demonstrated in fig1 and 13 , wherein fig1 highlights the inner ring 40 , which extends up to radial coordinate x 1 and requires more filtering to avoid artefacts in this region . the outer ring 42 is depicted in fig1 and requires less filtering / smoothing . after a filtering step e . g . with fixed size kernel has been applied to dataset 28 , the original and the filtered datasets may be combined using the weighting function w ( x ) shown in fig1 . this weighting function has a plateau of value 1 near the centre ( x = 0 ), and another plateau at a lower value in the area of the outer ring 42 . at the boundary , i . e . at the radial coordinate x 1 , a linear ramp is provided between the two plateaus . alternatively , the weighting function may be a step function between the central area 40 and the outer ring 42 . fig1 shows a preferred embodiment of the apparatus according to the invention . the invention can be embodied in an ultrasound machine 50 comprising a probe 10 , such as the tee probe shown in fig1 . the probe is connected to the control unit 51 of the ultrasound machine 50 , which comprises a processing unit 52 and a data storage unit 53 which are adapted to perform the method according to the invention . in addition , the processing unit 52 may be connected to a keyboard 54 and optionally a mouse or other cursor control device ( not shown ). furthermore , a screen 55 for viewing the output image dataset is also provided . alternatively , the method of the invention may be carried out offline on any type of computer system , which need not be part of an ultrasound machine . finally , fig1 and 17 demonstrate the effect of the filtering method of the invention to reduce near - field artefacts , while keeping far - field detail . fig1 is a surface - rendered image generated from a 3d ultrasound dataset of the heart . the 3d input dataset was generated by panning an ultrasound probe having a fan - shaped field - of - view , so that the input image dataset has a smaller pixel size in the near - field and a larger pixel - size in the far - field . fig1 shows the image rendered from the original raw data . since no filtering has been applied , blocky artefacts are visible in the near - field . on the same raw image dataset , the filtering method according to the invention was applied . a 5 × 5 kernel was used in the azimuthal - elevation direction , and a weighting function with a quadratic fall off from the near - to the far - field was used , due to the quadratic kernel n × n . from the thus computed output image dataset , a rendered image was generated using the same rendering algorithm as for fig1 . as can be seen on fig1 , a good filtering was achieved in the near - field , while keeping detail in the far - field .