1 Introduction
In recent years, 3D ultrasound (US) imaging has gained interest [1, 2] both due to the greater availability of 3D ultrasound systems (tracked, mechanical driven or matrix arrays)[3, 4]
as well as to the improvements in computational power allowing for faster and online processing of the acquisitions. To date, the most common way to obtain a 3D US volume remains the interpolation of 2D images into a 3D volume
[5]. We focus here on tracked 3D freehand ultrasound, a technique that yields 3D US volumes by first acquiring 2D images along with their position and orientation and then fusing them into a volumetric regular grid through a process known as 3D reconstruction or compounding [5]. Up to now, conventional compounding techniques (e.g. forward, backward or functional interpolation) have neglected an essential property of ultrasound images, namely, its directiondependent nature. It is known that imaging a point from different directions may lead to different intensities. This directiondependency is caused by the different attenuation paths followed by the reflected signals, as well as by the dependency of the reflected intensity on the tissue’s impedance and orientations as defined by Snell’s law. This directional behavior in part explains the information loss resultant from current 3D reconstruction methods[6] and is the main reason for restricted acquisition protocols (e.g. straight probe motion) recommended in practice. It is to be noted that such protocols modify the physician’s common practice based on image feedback and interactive motion of the probe.In this work, we introduce Computational Sonography (CS) as a novel paradigm for the 3D reconstruction and representation of tracked freehand 3D US that opens new possibilities for visualizing, interacting and processing volumetric data. In more detail, we introduce two major changes to current reconstruction paradigms (cf. Fig. 5). First, we consider and preserve the directional dependency of US in the 3D reconstruction. And second, we enable each element in the 3D volume to store a complex model instead of just a single scalar value. We argue that with such complex models it becomes possible to retain the directional information previously discarded and make it vailable for further use. We want to point out that the “Computational” term in CS, is an analogy to the concept of Computational Photography [7], and stands for the fact that once such a more complex volumetric representation has been stored, we deploy computational algorithms in order to extract and display the relevant information on demand.
The goal of this paper is to demonstrate how the concept of Computational Sonography can be translated into new directionpreserving compounding methods. In practice, we present two methods that reconstruct the amount of reflected signal expected from every direction at each point in a volume. An important design specification is the model to represent these values, where desired properties are the directionalitypreservation but also compactness, and fixedlength per voxel. We study two representation models complying with the requirements: tensors, as initially explored in a conference paper
[8], and a novel spherical representation. In summary, our contributions to 3D sonography are:
A reconsideration of tacit assumptions in the process of fusing multiple US views regarding i) the directionality of US, and ii) the reconstruction of scalar volumes.

The introduction of two models to represent multidirectional data, namely tensorCS and sphericalCS.

The analysis of two directionpreserving compounding algorithms and their comparison to classical compounding.

The demonstration of key capabilities of CS.
Finally, to show the practical value, we quantitatively evaluate the fidelity of the CS reconstructions (vs. information loss). The qualitative and quantitative evaluations are performed for both phantom and invivo data in comparison to stateoftheart compounding methods.
This work follows an initial conference publication [8] but includes the following extensions: in addition to the tensor model we propose a more expressive spherical representation, which encodes the directionality of ultrasoundinformation with minimal informationloss. We also further analyse indepth and quantitatively compare the two models against conventional compounding methods for both phantom and invivo acquisitions. In particular, we have extended the evaluation of the fidelity from 15 invivo and invitro cases to 40 cases.
1.1 Related Work
Only few methods have considered the directionality information in the processing 3D US volumes. Efforts in our group include a method to solve quadratic optimization problems in freehand 3D US in a compoundingfree fashion [9], as well as the design of directional dependent filters to vessels in 3D US images [10]. For 3D multipleview US reconstruction, Schulte zu Berge et al. [11] use the orientation information to cluster 2D slices, compound one volume per cluster and fuse the clusters. Different to CS, this method reconstructs only a scalar field, the directional information is lost for further processing.
Towards reconstructing more complex models per voxel, Klein et al. [12]
cluster the samples in a voxel according to their viewing angles but combine them into a volume of mixture models of Nakagami distributions, resulting in a vector containing the distribution parameters per voxel. However, the directionality information is also lost after reconstruction. Recently, Lee
et al. [13] and Papadacii et al. [14] proposed tensorialbased reconstructions (respectively Elastic Tensor Imaging (ETI) and Backscatter Tensor Imaging (BTI)) of US images from multiple view angles. Based on shear wave imaging, ETI measures the anisotropy of tissue elasticity and represents it with a tensor. In BTI, the local spatial coherence of the US signals observed for different orientations and focal regions was used to reconstruct 3D tensor volumes. Although these methods also recover tensor fields, their focus is on recovering the fiber orientation of the underlying tissues. Additionally, they both rely on a specialized hardware setup and a fixed scanning protocol. In contrast, our approach focuses on information preservation of the full acquisition and can be seamlessly integrated into the clinical workflow as well as applied directly to clinical 3D freehand US acquisitions. Finally, beyond tensor fields, we have also proposed a new representation for 3D US, SphericalCS.Representing US directional dependent information is similar to the problem of describing diffusion patterns in Diffusion MRI. In this field, both tensors (Diffusion Tensor ImagingDTI[15]) and spherical representations (Qball[16]) have been studied. However, apart from the representation model, the methods to fit and visualize tensors and qballs are specialized to the physics of MRI acquisition and therefore cannot be applied in a straightforward manner to 3D US data.
In conclusion and to the best of our knowledge, this is the first orientation preserving reconstruction method for 3D ultrasound. The proposed CSmodels bear some similarity to those used in diffusion MRI, but their computation from ultrasound data is new.
2 Computational Sonography Concept
We investigate the problem of building a directiondependent 3D US volumetric reconstruction from a set of tracked 2D US images consisting of a set of tracked US samples^{1}^{1}1US images are composed of rays, each formed by samples resulting from measuring the reflected signal at regular times.. We denote the set of acquired samples and perform the reconstruction on a 3D volumetric grid in a lattice , with . We can then decompose the reconstruction problem in two tasks:

Finding the relevant subset of samples contributing to the reconstruction of each voxel .

Fitting the chosen samples to the orientationdependent model stored in each voxel .
While we follow the strategy of [17] for the first stage (cf. subsec. 3.3), we focus hereafter on the reconstruction task. The problem can be formalized as finding a reconstruction function mapping from the set of measured samples to a model for every voxel in ,
(1) 
Let each US sample be characterized by a tuple , where is the measured reflected signal, and and stand respectively for the sample’s position and acquisition direction. In conventional compounding methods, maps the samples to a scalar value . Accordingly, the mean reconstruction can be formalized as
(2) 
Other common scalar reconstructions include the median, kNearest Neighbor or Inverse Distance Weighted compounding methods.
With computational sonography, we propose two new reconstruction functions , which fuse samples while preserving ultrasound’s directionality; i.e. they capture the fact that the reconstruction may show different intensity values for different directions. The two proposed methods, one tensorbased and another sphericalbased , differ in the fusion as well in the final model used for compact, but direction preserving, representation of the reconstruction, as detailed in sec. 3.
The result of the two CS reconstruction procedures is a compact representation of the full acquisition, which preserves the directional nature of US for visualization, as well as other possible common postprocessing tasks.
3 Methodology
In order to design a directionpreserving reconstruction function , we propose to store a representation model instead of a scalar value in each element of a target lattice. We present two reconstruction functions and representation models. The first, initially introduced in [8], is a tensorbased approach, Tensor Computational Sonography (TensorCS). The second, is a novel approach relying on the concept of spherical grids (SphericalCS).
3.1 TensorModels
For TensorCS, we choose to fill the output volumetric grid with tensors, that is we store a tensor model in each voxel. The reconstruction function of TensorCS is implicitly defined as:
(3) 
where the tensor is assumed to be symmetric. The direction preservation is achieved by means of the constraint , which states that tensor should compactly encode the intensity value observed from the given direction . In practice, a least squares approach is employed to find the values of the tensor , where we minimize
(4) 
As shall be symmetric, at least six sampling points determine the tensor coefficients. In practice, we use typically between 12 and 500 samples, as this showed to provide best results in our experiments. We solve the least squares problem using Cholesky decompositions implemented on a GPU.
The main advantage of the tensor model is that it allows for a direction dependent modeling of the ultrasound signal while requiring only six real values to be stored per voxel, making TensorCS very suitable for parallel processing. The tensor model is also suited for an interactive orientationdependent visualization (see subsec. 3.4). Finally, tensor fields can be further postprocessed by means of tensorbased regularization and visualization methods, such as those used in DTI [18, 19, 20]. On the other hand, the tensors’ compactness comes at the price of resolution loss and some simplifications. For instance, the tensor modeling assumes that signals are point symmetric. While the symmetry is irrelevant for the most common freehand acquisitions, it becomes a problem for acquisitions from all orientations, and in particular, in the presence of acoustic shadowing. This phenomenon is illustrated in Fig. a and Fig. b.
3.2 Spherical Models
To overcome some of the limitations of TensorCS regarding the resolution and enforced symmetry, we propose a new representation for 3D US data. More precisely, for each voxel we model the ultrasound signal reflected from all directions as a continuous real function defined on the unitary sphere . This model allows a unique intensity value for every direction , or equivalently, . As a result, the SphericalCS representation consists of one spherical function per element of the lattice , that is, . Recall that the samples relevant to voxel are defined by , i.e. by their intensity , position and imaging direction . We then try to approximate the spherical function from the samples’ tuples . Although one could store every sample relevant to each voxel to approximate with the highest fidelity, the resulting representation would be very inefficient, due to a varying number of samples. Instead, we choose as domain a spherical grid, computing a discrete approximation of . The domain is defined as a partitioning the sphere into segments , such that . Then, we approximate a scalar value per cell. Finally, we collect the values from the different cells to store a model in each voxel. Thereby, the sought reconstruction function is:
(5) 
In practice, each entry of the vector is computed only from the samples falling inside the corresponding cell . First we determine the subset of samples associated to each segment . Then, the value assigned to the cell is obtained from the samples intensity values through one interpolation schemes of
(6)  
(7) 
where stands for the number of samples in segment . They are derived from classical compounding methods, but applied only to the associated sample subset
This discretization scheme permits adjusting the desired resolution of the representation. In this way, we achieve a compact, yet accurate representation of the acquisition allowing for subsequent processing or extraction of information according to the needs of the user. The proposed models are also well capable of representing arbitrary functions including nonpointsymmetric data (cf. Fig. 10).
A challenge with the spherical representation arises from the discretization of the sphere, which is itself an open subject of research [21]. We are interested in a discretization that: i) leads to a regular partition of the area on the sphere, and ii) enables computational efficient mappings from a directions to the respective segment and backwards, as these mappings will be needed for the reconstruction as well as for later access, processing or visualization.
Next, we analyse and discuss three possible discretization approaches, as well as their mappings. To this end, we first define a grid on the unit sphere as a set of points , where each point is described by its spherical coordinates, namely the azimuthal and elevation angles, i.e. . Then, the Voronoi cells of the grid points inherently define a partition of the sphere.
3.2.1 Latitudelongitude grids
Given a sphere parametrized with spherical coordinates, the latitudelongitude discretization is simply obtained through an equidistant sampling of the azimuth and elevation angles. As a result, it is trivial to map direction vectors to their respective point on the spherical grid by converting Euclidean coordinates to spherical coordinates and visceversa. More precisely, the coordinates of the points on a latitudelongitude grid with a resolution of radians are given in spherical coordinates by
(8)  
(9)  
(10) 
As it is apparent in Fig. a, the inconvenience of latitudelongitude grids is that the density of points varies significantly across the surface of the sphere, which results in cells of variable area, and is undesired for computational sonography.
3.2.2 Icosahedral geodesic grids
Geodesic grids are built by iteratively subdividing a polyhedron. It is common to use a icosahedron as a starting point. At every iteration, each polygon of the current polyhedron is further divided in cells, and projected onto the sphere. This subdivision is continued until the desired resolution is reached. Following this discretization, the obtained partitions have nearequal area. For a detailed description we refer to the work of Thuburn [22]. Fig. b shows a geodesic grid based on an icosahedron subdivision. Similar to the latitudelongitude grids, the spherical coordinates of the gridpoints are known by construction and the mapping to their represented directions is trivial. However, no constanttime mapping of arbitrary direction vectors to the corresponding gridpoint (or directional segment) exists for these grids [23]. Since measurements can originate from arbitrary directions, icosahedral geodesic grids are not practical from a computational point of view.
3.2.3 Fibonacci geodesic grids
The third type of discretization we studied here are spherical Fibonacci grids, which are generated by drawing points along a spiral on the surface of the sphere. The construction rules describe each gridpoint in terms of the golden ratio [24], as follows:
(11) 
Fig. c shows a spherical Fibonacci grid with 42 gridpoints. Mapping a gridpoint to a direction vector is equivalent to converting spherical to Euclidean coordinates. Most importantly, however, the inverse mapping recently proposed by Keinert et al. [23] provides a mapping from any direction vector to the nearest gridpoint or directional segment with constant runtime. We refer the reader to the work for the derivation of the inverse mapping. This property makes Fibonacci geodesic grids the most suited for computational sonography.
3.3 Sample Selection
As discussed in sec. 2, the compounding process is composed of two stages. So far we focused on the reconstruction step and assumed the subset of measurements relevant to fit one voxel model is known. This subset is determined using a sample selection strategy initially described in [8], and inspired by [17]. For completeness, we describe the process next.
The task is to select all samples which contribute to the reconstruction of a specific voxel, centered at coordinates . Practically, we select a sample if the voxel center lies inside the region of influence of the ultrasound sample. This region is defined as an ellipsoid centered around the sample’s position and with one of its axis aligned to the sample’s direction . Formally, if:
(12) 
where represent the half of the length of the ellipsoid’s axes in each direction.
Fig. 16 shows the sample selection process exemplarily. Based on (12), multiple samples of each scan line could be selected depending on the proximity of a ray to the target position . As the main interest for the reconstruction is to collect complementary information, only the nearest sample of each ray is retained in .
3.4 Computational Aspects of CS
After the samples have been selected, and a model (either tensor or discretized spherical function) has been computed per voxel, the 3D reconstruction process is finished, and the result is a 3D lattice of models. In this section we discuss the “computational" aspect of CS, where we will make use of algorithms to extract the relevant information from the lattice both ondemand and according to the application. Within the scope of this paper we restrict the algorithms to the extraction of the information needed for visualization.
3.4.1 Tensor visualization
From the TensorCS reconstruction, it is possible to extract directional images, i.e. images of the reflected ultrasound signal from a given direction. The directional intensity for a voxel from direction , with
, can be estimated from the tensors with the second order approximation:
(13) 
Thereby, (13) allows us to interactively visualize the image as if taken from an arbitrary viewpoint, even if not previously imaged. Such interactive directiondependent visualization can serve as a basis for realistic simulations for training.
Secondly, from the TensorCS we can also retrieve a variety of 3D scalar representations suitable for simple 3D renderings. The first option is to compute the absolute value of the tensor traces and use them as the intensities. In this way the intensity at voxel becomes
. The tensor trace amounts to display the accumulated signal intensity from the three major components (eigenvectors) of the tensor. A second possibility is to extract the intensity from the tensor’s dominant orientation corresponding to the largest eigenvalue of the tensor
.3.4.2 Visualization of discretized spherical functions
In the case of the SphericalCS, and in particular for the geodesic grids, we studied the visualization of two quantities: the mean and the maximum of the intensity values stored in the different cells of the geodesic grids.
(14)  
(15) 
These values are suitable for showing respectively, the most representative intensity in each voxel and the intensity from the direction with highest response. Examples for these visualizations can be seen in subsec. 4.4. These choices are motivated to maximize the speed of the visualization task. It is worth noting though that the representation is open to more complex reconstruction approaches.
4 Experimental Validation
To evaluate the different reconstruction capabilities of Computational Sonography we present next a series of experiments on both phantom and invivo datasets (subsec. 4.1). First, we perform a quantitative evaluation of the CS representation quality in terms of information preservation and in comparison to classical compounding (subsec. 4.2). We then investigate the robustness of the representation within a voxel to increasing variations of the intensity values, or to an increasing coverage of the acquisition directions (subsec. 4.3). We also visually illustrate the direct use of the CS representation to create scalar 3D volumes (subsec. 4.4). Finally, we present exemplary uses of the additional information present in CS volumes in subsec. 4.5. The evaluation is based on a comparison of the following spatial compounding methods:

Mean compounding.

TensorCS.

SphericalCS with latitudelongitude grid.

SphericalCS with spherical Fibonacci grid.
For all the SphericalCS reconstructions we used the mean of sample intensities in each cell (6). For the TensorCS one constraint not enforced so far is the expected positivedefiniteness of the tensors (), arising from the the fact that intensities are by nature positive.
4.1 Invitro and Invivo Data Acquisition
TypeSeq. No.  1  2  3  4  5  6  7  8  9  10  11  12  13  14  15  Total 

Vessel phantom  125  217  556  633  725  1166  149  269  100  561  592  1000  6093  
Carotid  160  50  292  336  437  479  224  379  25  197  209  187  204  3179  
Femoral  89  377  312  78  209  188  306  81  31  249  91  232  204  90  187  2724 
We acquired a number of phantom and invivo datasets using a 3DUltrasound system consisting of an Ultrasonix SonixMDP machine with a linear transducer and a NDI Polaris Vicra optical tracking system. Every sweep is composed of several Bmode images acquired in a sequence while smoothly moving the probe. All acquired 3DUS sweeps differ in the number of images and in the span of directions covered. We followed a protocol where we gradually varied from sweeps taken for a single direction (as commonly done in today’s 3D US freehand protocols) to sweeps densely covering the sphere.
In practice, we acquired 12 sweeps of a static latex vessel phantom (cf. Fig. 17) and 28 invivo sweeps, including 13 vascular scans (carotid arteries) and 15 scans in an osteological setting (femoral bones). The number and size of the different acquisition sequences are detailed in Table 1 while the acquisition parameters are summarized in Table 2.
Parameter  Value 

US probe  L145/38 Linear 
Frequency  6.6MHz 
Depth carotid  4.0cm 
Depth femoral  5.5cm 
Depth vessel phantom  4.0cm 
Sample selection  (12) 
Reconstruction resolution  0.5mm spacing 
4.2 Representation Quality
Using all invitro and invivo datasets above we evaluate the representation capabilities of CS. First, we evaluate the representation quality in terms of how well a reconstructed volume preserves the information from the original ultrasound images, which is standard evaluation criterion for 3D reconstruction methods [5]. In particular, we focus on quantifying the preservation of the directional intensity, i.e. the intensity observed from different directions. To this end, and in contrast to the evaluation of classical compounding methods based on restricted scanning protocols (i.e. parallel images), our datasets contain information from various probe directions.
To quantify the information preservation, we reproject the volumes at the positions and directions from where the original US images were acquired. For a volume reconstructed from a sequence including samples , we measure the Mean Squared Error (MSE) between the input and the reprojected intensity values of every sample as follows:
(16) 
where is the intensity of sample and is the intensity computed from the reprojection of the compounded volume at the position and direction of the original sample . The error is averaged over all input samples. For all experiments, 3D volumes were reconstructed with a 0.5mm spatial resolution. For the latitudelongitude grid a resolution of was used, whereas for the Fibonaccigeodesic reconstruction the grid was built with cells. For scalar volumes, the reprojection simply corresponds to interpolation. In the case of tensorCS, we reproject using (13), whereas for sphericalCS we use the value of intensity stored in the associated grid cell. Examples of the reprojections for each method are shown in image Fig. 30. It can be seen that the 2D projections from computationalsonography volumes better preserve the information from the original Bmode images as CS avoids averaging information from different directions. Although the tensorCS recovers appropriate intensity values for large portions of the image, the optimization fails in voxels where the information exceeds the lowcomplexity of the second order model or the symmetry. This is reflected as white regions in the images.
Mean  TensorCS  LatLongCS  GeodesicCS  

Carotid  0.024 0.034  0.024 0.092  0.021 0.031  0.014 0.028 
Femur  0.028 0.039  0.038 0.157  0.022 0.034  0.015 0.030 
Invitro  0.025 0.031  0.061 0.242  0.020 0.029  0.008 0.024 
Overall  0.026 0.036  0.040 0.170  0.021 0.032  0.013 0.028 
We report in Table 3
the mean and standard deviation of the representation error in (
16) for all volumes computed from the invivo and invitro acquisitions. For a fair comparison with the tensorCS, we considered only the representation error of voxels for which a tensor could be estimated. The results indicate that representing the Bmode data by tensors leads to errors comparable to mean compounding, which is in accordance to the results in our prior work [8]. SphericalCS based on Latitudelongitude grids does not substantially reduce the representation error, whereas geodesicbased sphericalCS results in the lowest average errors for all datasets. Fig. 31 shows violin plots of the representation errors found in a carotid acquisition (Seq. 5, 437 frames, ROI dir.var. 0.0257 0.0061) and emphasizes the similarity of error distributions between the presented representations.4.3 Robustness to Intensity Variability and Spherical Coverage
In the following experiments, we investigate the robustness of the compared methods to two potential sources of inconsistency within the samples of a voxel: intensity variability and directional coverage.
To evaluate the robustness against the intensityspan
, we calculate the representation error for an increasing variance of intensity values associated to a voxel. Given intensities normalized to
, we computed the variance of the intensities for each voxel , and used a histogram to group voxels with similar variance. Figures (a)a and (b)b show the representation error for the different quantized levels of intensityvariations. It can be clearly seen that all Computational Sonography methods results in significantly lower errors, especially for larger values of intensity variance. Moreover for SphericalCS the error is stable.We also analyze robustness to the sphericalspan, that is, how the variance in the span of directions covered by the acquisition affects the representation error. We repeat the procedure above, but quantizing the spherical variance , as defined by Mardia and Jupp [25], where is the so called mean resultant length, which depends on the clustering of the vectors around the the mean direction. Note that is in
, corresponding to acquisitions going from a single direction to uniformly distributed over the sphere. We then estimate the representation error for the different levels of spherical variance. Figures
(c)c and (d)d shows how an increasing variance of directions affects the tensor representation the most, which is explained by the fact that the symmetry assumption is more likely broken in voxels compounding samples with high spherical variance. For the meancompounding and sphericalCS the increase of spherical spans does not seem to have a noticeable impact on the reconstruction error.The two graphs provide an experimental evidence to the fact that classical compounding approaches result in reconstruction errors when applied to unconstrained freehand acquisitions. The reason for this is not directly the increased spherical variance but rather the ultrasound physics causing the same structure to have large intensity variations in different views. With regard to the two different CS representations, the tensorCS suffers from a limited 2nd order model complexity, which reflects in a poor performance under large intensity variations On the other side, the strongest advantage of sphericalCS comes for acquisitions with heterogeneous ultrasound information (i.e. strong intensity variations) or high spherical variance. In fact, the error of sphericalCS reconstructions decreases for larger spherical variances, better achieving the goal of preserving the directional information, although at the cost of increased model complexity. These results further confirm computational sonography as a viable alternative to traditional reconstruction methods, allowing for high quality 3D freehand reconstructions from arbitrary scanning trajectories instead of restricted scanning protocols.
4.4 Qualitative comparison
Exemplary results of the different reconstruction methods are shown in Fig. 62. They include axial and lateral slices for each of the sequence types: femoral, carotid and vessel phantom. For tensorCS, we show the absolute value of the tensor trace. For sphericalCS, we use two forms of visualization: first, the intensities reflecting the mean of the cell values excluding empty cells, and second, the colorcoded direction of the cell with the highest value. The color coding maps each vector in the volume to a unique color (cf. Fig. 37)^{2}^{2}2In practice, we first estimate the peracquisition main direction by means of averaging the direction vectors in the volume. The colorcoding then shows deviations from the main direction with increasing saturation, the hue is determined by the angle on the plane perpendicular to the main direction. Because the map considers 3D direction vectors, similar colors in different views stand for similar maximum directions.
The images reveal first that TensorCS has an unstable behavior in unconstrained large scale evaluations. Given its reduced representation complexity it is prone to fail when the data from different directions is of higher order than our 2nd order model (the intensities vary significantly across different directions in a complex pattern). With the current implementation this effect is translated into a failure of the tensor fitting, shown as white voxels in the image. This effect seems to be located in boundary areas and locations where numerical instabilities are evoked by low intensities measured from a high number of directions, as can be seen in the phantom views. Despite this behaviour, it is apparent that whenever the tensor calculation succeeds the details of soft tissues are increased. Moreover, the tissue interfaces, e.g. in the two views of the carotid sequence, are best defined in TensorCS. In the case of SphericalCS, the cellmean visualization filters artefacts coming from directions with secondary reflections. This can be clearly seen for the femur lateral view: while the mean compounding features reflections behind the femoral interface, these reflections do not appear in the sphericalCS mean reconstruction. Also, a considerable blur reduction is observed when comparing sphericalCS to the mean compounding, in particular for the two views of the carotid and the lateral vessel sequence. The color visualization of the maximum SphericalCS cells provides information about the interfaces’ normals, as can be seen in the femoral lateral view but more clearly as the normal of the carotid changes in the carotid lateral view or in the vessel phantom.
4.5 Free view point evaluation
The directional information preserved in the CS models can be used to estimate intensities from arbitrary directions. This is achieved by evaluating (13) for tensorCS with a chosen direction. Fig. 65 shows this for a carotid scan (Seq. 5), where the left is extracted from the original viewpoint and the image resulting from another direction on the right. It can be seen that the contrast between the interfaces and the tissue intensities is changed between the two directions.
5 Discussion
The qualitative results in combination with the quantitative evaluation show that geodesic spherical CS provides the best representation for arbitrary USsweeps, while complying with the requirements of data fidelity and constant storage. The mappings from samples’s directions to the sphere are slightly more complex than for TensorCS, but this is compensated by the fact that as opposed to TensorCS, SphericalCS does not rely on an optimization procedure to reconstruct the cellvalues for a given voxel. As this optimization is based on all samples in a voxel, it can only be performed in a twostep reconstruction as we have presented it. The sphericalCS on the other hand could be implemented in an iterative reconstruction process, thus allowing for runtimes comparable to state of the art scalar reconstructions. In addition, the current implementation of TensorCS has shown an unstable behaviour for general acquisitions with wide spherical variation. We believe that part of the instability could be solved by using more sophisticated optimization algorithms resulting in additional computational cost. However, the inherent limitations of the 2nd order symmetric model would still remain.
The concept of Computational Sonography opens new visualization paths for 3D US, relying on CS ability to store and then compute values on demand. In addition to the visualizations in Fig. 62 and according to the application, one may want to display the most meaningful cell for each voxel, to ensure faithfulness to the original measures. Online interactive visualization methods can be used to visualize the true directionaldependency of the ultrasound volume even after reconstruction and as such can serve as a basis for realistic simulations for training. Finally, we may also use the directional information to guide online acquisitions towards a viewpoint where the structure is most visible or to create surfaced visualization of interfaces.
Beyond visualization, Computational Sonography aims at compactly and reliably encoding all the information available during a freehand acquisition in order to make it available for later processing. For instance, the CS concept is also useful for segmentation where the direction of the maximum cell gives hints about the presence of surfaces. Also for registration, CS allows us enables to compute the better similarity functions between a 2D slice and the reconstructed 3D volume, when accounting for the directional dependency of the images.
6 Conclusion
We showed the performance of several methods for the representation of the directional dependent information found in freehand US, using the conventional mean compounding as a baseline. Two storage models were proposed. A tensorCS model, which we initially proposed in [8] and a novel approach based on spherical grids. We illustrated the adaption capabilities of the two representations, showing that the information can be extracted on demand according to the application. Finally, an exhaustive quantitative evaluation of the representation error resulted in a favorable performance of the spherical grids over the tensors and mean compounding methods. These results serve as evidence to our claim that with our new representation freehand acquisitions do not any longer need to rely on linear protocols.
Future research includes exploring other applications that require more complex algorithms to extract the relevant information, comparable to compressed sensing approaches. Additionally, one could devise methods using nonlocal information, for example to extract surfaces of tissue interfaces.
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