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Strain and strain rate parametric imaging. A new method for post processing to 3-/4-dimensional images from three standard apical planes. Preliminary data on feasibility, artefact and regional dyssynergy visualisation
© Støylen et al; licensee BioMed Central Ltd. 2003
- Received: 25 June 2003
- Accepted: 25 August 2003
- Published: 25 August 2003
We describe a method for 3-/4D reconstruction of tissue Doppler data from three standard apical planes, post processing to derived data of strain rate / strain and parametric colour imaging of the data. The data can be displayed as M-mode arrays from all six walls, Bull's eye projection and a 3D surface figure that can be scrolled and rotated. Numerical data and waveforms can be re-extracted.
Feasibility was tested by Strain Rate Imaging in 6 normal subjects and 6 patients with acute myocardial infarction. Reverberation artefacts and dyssynergy was identified by colour images. End systolic strain, peak systolic and mid systolic strain rate were measured.
Infarcts were visualised in all patients by colour imaging of mid systolic strain rate, end systolic strain and post systolic shortening by strain rate. Reverberation artefacts were visible in 3 of 6 normals, and 2 of 6 patients, and were identified both on bull's eye and M-mode display, but influenced quantitative measurement. Peak systolic strain rate was in controls minimum -1.11, maximum -0.89 and in patients minimum -1.66, maximum 0.02 (p = 0.04). Mid systolic strain rate and end systolic strain did not separate the groups significantly.
3-/4D reconstruction and colour display is feasible, allowing quick visual identification of infarcts and artefacts, as well as extension of area of post systolic shortening. Strain rate is better suited to colour parametric display than strain.
- Systolic Strain
- Colour Display
- Strain Rate Imaging
- Minimum Strain Rate
- Longitudinal Strain Rate
Colour tissue Doppler  samples tissue velocities, nearly simultaneously from all pixels in the 2D picture. Data are quantitative, and can be displayed as curves or numbers. Quantitative analysis in all segments of the left ventricle, however, is time-consuming. Colour display, or parametric imaging, reduces the displayed data to semi-quantitative information, allowing a quick visual assessment of functional data over a larger area. In addition, this display allows new measurements, such as area and propagation velocity.
Velocity data, being simultaneous and digital, can be post processed to derived parameters. Numerical integration of velocity over time, results in displacement curves, reproducing the motion curves from M-mode. Both velocity and displacement are imaging of local motion. Tissue tracking, an imaging modality developed at the Karolinska Institutet in Stockholm , displays displacement by bands of different colours (fig. 1b). In this modality global function is visualised by the annular displacement, which is related to EF [3–5], local motion by the colour, regional differences by the symmetry and deformation by the width of the bands. Tissue tracking is thus the only parametric imaging modality that displays both motion and deformation at the same time.
Strain rate imaging, a modality developed at the Norwegian University of Science and Technology in Trondheim , is a spatial derivation of velocity: SR = ((v(x) - v(x + Δx)) / Δx in the direction of the ultrasound beam. This algorithm subtracts regional velocities due to translation and tethering, giving in the rate of regional deformation – strain rate, an index of regional contractility . Clinical findings of reduced regional function in myocardial infarction have been validated against 2D echocardiography  and coronary angiography . In strain rate, yellow to orange represents shortening, cyan to blue lengthening (fig 1c). The numerical derivation results in a less favourable signal – to – noise ratio in strain rate than in velocity [9, 10], resulting in increased random noise. Strain – local deformation – can be measured by temporal integration of strain rate. Strain measurements by this method have been validated against ultrasonomicrometry  and MR . The integration process eliminates most of the random noise, but strain is still sensitive to non-random noise, and in addition prone to baseline drift. Both strain and strain rate are deformation imaging, as opposed to motion, and are sensitive for regional ischemia. They are translation and tethering independent, as opposed to motion, and will thus show the true location of regional dyssynergy [6, 8, 9]. On the other hand, experience with the method is necessary for discerning true pathology from artefacts due to dropouts and reverberations .
We present a method where the velocity data from 3 standard apical imaging planes are acquired, integrating spatial information and the temporal sequence results in a four-dimensional data set. From this, curved M-modes, bull's eye views and a 3-dimensional surface figure can be generated. The four-dimensional nature of the data results in the ability of the display to be scrolled through the heart cycle. The method also includes the post processing ability to derive any secondary parameter such as motion, strain rate and strain and the ability to re-extract numerical data as curves or numbers. The basic principles are illustrated in fig. 2.
The aim of this study was to evaluate the feasibility of this imaging modality in a clinical setting, to display regional dyssynergy and artefacts and to see if regions of maximum and minimum contractility could be identified for quantitative measurement.
Six patients with first, acute myocardial infarction, two inferior, four anterior, were examined. The patients participated in a prospective study of strain rate in myocardial infarction, and examinations were done during the first two days. Six normal subjects were examined as controls. These participated in a stress echo study, selected on the basis of a normal coronary angiography done for clinical reasons, and all had normal resting echocardiogram by conventional criteria. Resting cine-loops were used for the present study. All studies were approved by the regional ethical committee, and written consent was obtained. The patients were included consecutively, and no patients excluded for poor image quality.
Echocardiography and post processing
Echocardiography was performed with a GE Vingmed Vivid 7 scanner (GE Vingmed Ultrasound, Horten, Norway), with tissue Doppler acquired at a frame rate of 100 – 150 FPS. Single cine-loops from each standard apical plane – 4-chamber, 2-chamber and apical long axis – were transferred to a PC computer for post processing in experimental software (GcMat, GE Vingmed Ultrasound), programmed in Matlab (MathWorks inc., USA).
First, a curved M-mode was drawn in each of the three planes by placing seven points and curve drawn automatically by spline interpolation. Velocity data were sampled along the line. Temporal smoothing over 3 frames, and spatial smoothing over 3 pixels was applied automatically. Strain rate and strain was processed with an offset length of 12 mm. Adjustment for unequal RR-intervals was achieved by resampling of 200 points of the ECG curve, and the velocity data interpolated to fit the ECG curves. For rotational angle between planes, an assumption of 60° between the planes (being the average separation) was applied in the program, and velocity data were interpolated by cubic spline by a method described previously . Data was then displayed as bull's eye (fig. 1, top), curved M-mode array (fig. 1, middle) and 3-dimensional surface grid (fig. 1, bottom). In 3D display, the imaging planes and the aortic annulus were added for orientation. The whole of this post processing took 2 – 3 minutes.
The points of maximum a minimum strain rate in mid systole were identified visually, and mid systolic and peak systolic strain rate as well as end systolic strain were measured in these points. Care was taken not to avoid the areas of reverberations or dropouts, as the object was to evaluate the impact of artefacts in this feasibility study. The differences between maximum and minimum strain rate was calculated.
Infarcts were visible as inhomogeneous mid systolic strain rate and end systolic strain; both parameters located the infarcts equally well. Post-systolic shortening was evident in early diastolic strain rate in all patients. The area with post-systolic shortening was visually larger than the area of dys- to hypokinesia in mid systole. Post-systolic shortening could not be identified reliably in colour imaging, even when present in the curves. Reverberations were present in one patient, basal artefact in one, both in strain rate and strain.
Average maximum and minimum mid systolic and peak systolic strain rate and end systolic strain.
Mid syst. strain rate (s-1)
Peak syst. strain rate (s-1)
End syst. strain (%)
Neither the trend toward hypokinesia of the infarcted areas by mid systolic strain rate and end-systolic strain, nor the trend toward hyperkinesia in non-infarcted segments in patients, was significant. Reverberations could not be reliably identified from curve analysis alone (fig 5).
Basic limitations of strain rate and strain imaging
It must be emphasised that the present study is about the method of 3-dimensional reconstruction and display, while the basic method of strain rate imaging is not different from that published previously [6–12, 14]. Strain rate imaging in itself is based on post processing, and has a number of inherent limitations.
The most basic limitation is the one-dimensional nature of the data. While deformation is three-dimensional (longitudinal, transverse and circumferential), the present tissue Doppler derived method measures only longitudinal deformation. However, heart muscle is generally assumed incompressible, which means conservation of volume. Hence, strain in three dimensions must balance (the basic equation in volume conservation: εx + εy + εz = 0), so in systole longitudinal plus circumferential shortening equals transverse thickening, and vice versa in diastole. Previous studies have shown similar diagnostic information by longitudinal strain rate and transverse wall thickening (which is the same as transverse strain) by ordinary echocardiography . This means that the longitudinal strain rate may be representative of the compound regional deformation. The practical value for clinical work of adding more deformation directions remains to be demonstrated.
Data represents only a line along the wall, but to achieve a high frame rate, the tissue Doppler beams have to be wide, and the lateral resolution is low anyway. This means that differential strain across the wall cannot be measured, however, this is true for all tissue Doppler derived data sets.
The one-dimensional nature of the data also results in angle distortion, when the ultrasound beam deviates from the direction of the motion. While velocity measurements decrease by the cosine of the insonation angle, strain rate decreases more, adding the transverse strain rate with the opposite value [6, 18]. This means that in areas where the angle between the wall and the ultrasound beam is more than 45 degrees, the strain rate values and hence, the colour will reverse [11, 18]. This problem is apparent in the apex and the base . However, all clinical studies of strain rate have found it possible to measure strain rate and strain in the apical segments, by measuring in the basal parts where the walls are more parallel to the ultrasound beams [8, 9, 12, 14, 19, 20]. As can be seen from fig. 4, the area of distortion is quite small, and during the last part of the study, blanking was applied to this area to avoid displaying misleading information. This limitation is more theoretical than practical.
Strain rate / Strain imaging are prone to reverberation artefacts, due to the basic algorithm. If one of the two velocities is measured as zero, the strain rate value will be inverted in part of the band. The 3D reconstruction does not add to this, but makes it visible, it is the identification, not the presence of artefacts that are influenced by this method.
Strain rate is sensitive to random noise, and smoothing of curves has to be applied, resulting in loss of temporal resolution. This might lead to underestimation of peak values, but in parametric imaging, being semi-quantitative, this is less important.
3D reconstruction is a method for quick visualisation of global distribution of deformation. Infarctions were visible as dyssynergy in the typical anteroapical or inferior locations, and colour display of mid systolic strain rate and end systolic strain were equally sensitive. Post-systolic shortening , was an additional finding in all infarcts. Thus, it represents an additional diagnostic criterion, facilitating visual diagnosis of acute infarction. The mechanism for the phenomenon is incompletely understood, but has been clearly documented as a marker of acute ischemia [21, 22]. It is interesting that the area of post-systolic shortening was larger than the area with hypo- to dyskinesia in all patients. The difference may show an additional area of partly ischemic myocardium, visualising the border zone at risk. This has been visible by curved M-mode alone in previous studies as well [8, 9], although not commented on. Post systolic shortening was visible in both strain rate and strain curves, but only in strain rate colour display.
Reverberations were easily identified. On M-mode they did show up as horizontal bands, on the bull's eye as ring shaped structures. The location, shape and temporal sequence differed from the typical vascular areas, and were thus discerned from true pathology, as shown in fig. 5. The artefacts were not so obvious from the curves alone. Basal artefacts are due to either angle distortion or inclusion of part of the mitral ring or pericardium during initial processing. It is important that in three out of six normal subjects, reverberation noise was visible. This high incidence of artefacts has not been reported previously, and may have important implications for the validity and accuracy of quantitative data in clinical practice.
By identifying regions with artefacts, the colour display may aid in avoiding quantitative measurements in those, and thus may also increase the specificity of quantitative strain rate / strain in clinical practice.
Strain rate vs. strain
Deformation imaging is better in identifying regions of dyssynergy than motion imaging, due to tethering effects. Velocities decrease from base to apex  while peak strain rate is the same in all three levels [10, 14, 19, 20].
Peak strain rate was the only quantitative parameter in this study able to identify the infarct areas, while neither mid systolic strain rate nor end systolic / peak strain could do this. This demonstrates a basic limitation of 3D reconstruction. In 3D display, all parts of the ventricle are (in theory), displayed at the same instance of the heart cycle. All previous works in strain rate, however, have concentrated on peak strain rate, but peak strain rate has not been demonstrated to be simultaneous in all parts of the ventricle. Thus, mid-systolic strain rate (or strain rate at any other point in systole) is not equivalent to peak strain rate. Systolic strain is maximal in end systole in normals, and peak strain thus simultaneous. In theory this should be an advantage in 3D reconstruction, but the presence of reverberations eliminated this.
There was no difference between strain and strain rate in the sensitivity to reverberations and dropouts (fig. 4). Smoothing of strain rate data results in reduction in measured peak values, but the temporal resolution is still better in colour strain rate than strain, due to the shifts between positive and negative values as evidenced by the visualisation of post-systolic shortening. Thus, strain rate is best suited to colour display.
Limitations and advantages of the processing method and 3D display
It may be considered a major limitation that the circumferential resolution is limited to three planes, and that most of the information displayed is interpolation. However, the main point of this method is about display of data. The 16-segment model of the left ventricle  is the basis of all analysis of regional function in clinical cardiology, wall motion scoring as well as quantitative measurement by velocity, displacement and strain rate / strain imaging. In all these modalities, measurements are taken as representative of a whole segment, resulting in a circumferential resolution of 60°. This is true of any Bull's eye plot, whether of wall motion score or numerical measurements. Reconstruction to Bull's eye alone, does not add more information, but rather displays it in a more accessible form. The method itself, however, is not limited to three planes, but more planes will be non-standard, however. With a gated rotational probe, the number of planes could be increased, but the post processing time would increase proportionately. However, doubling the number of planes, the whole post processing time would still be less than five minutes. The impact of the number of planes has been extensively studied previously . Three planes result in a correct average of volume and surface area, even in asymmetric ventricles. Variability is high, and decreases substantially with an increase to four planes, and little is gained by further increase. Concerning regional dyssynchrony, more planes would increase the exact delineation of the areas. This, however, applies to the whole concept of segmental analysis.
Using three planes, the average angular separation is 60°. This is in accordance with the standard representation of bull's eye, and was chosen for the visual display. The true position of the planes is closer to 0, 62 and 101°. These values could be implemented into the model at need, but previous analysis have shown that a variation of ± 15° results in less than 2% error , so the practical advantage of this correction is dubious.
The radial resolution (along the ultrasound beams) is a little more than strain offset length, around 15 mm; i.e. about twice as good as that of the 16-segment model.
Bull's eye and M-mode array could be obtained more directly, without the 3D reconstruction. The 3D surface model, however, incorporates the curvature data, which truly adds new information. The main advantage of this is giving an approximate representation of the true area. By subjective assessment, it was an additional help in differentiating infarct areas from basal artefacts in two patients, as it gave a more correct display of the shape and size of the infarct area. The method has further possibilities of true quantitative infarct area measurement, subject to validation studies.
The 3D surface model gives a more intuitive visualisation of the left ventricle, for clinicians unaccustomed to bull's eye or curved M-mode. However, this is also a potential source of pitfalls. The processing method will result in a grid, whether there are tissue data or dropouts. Appropriate blanking of the apical area should be applied to avoid interpretation of angle distortion in the apex as dyskinesia, but as already discussed; this region is far smaller than the extent of the apical segments. Finally, the representation still only shows longitudinal deformation, but as discussed previously this may reflect compound deformation in more than one dimension.
Combining data in a four-dimensional data set results in versatility. The method described may be seen as a method for storing data, where any desired derived parameter and display method can be extracted. Segmental waveform analysis, as well as any kind of parametric display modality as is M-mode, bull's eye and 3D surface is then available. As shown, each type of display contributes to the total information: Bull's eye by showing the whole of the surface, M-mode by showing space – time relations and 3D surface by displaying the anatomy of the surface and the correct area.
Compensation for HR variability during examination
With varying HR, the RR-interval will vary too. Systole and diastole varies differently, so a simple linear correction ("stretching") by the end points of the heart cycle is insufficient. A better approximation would be to correct systole and diastole separately, by identifying a point in end-systole, and then fit the two parts of the cycle separately – a three point correction. The present algorithm instead uses a 200-point correction by ECG, so the correction should in theory be correct down to 5 ms, but the use of ECG makes it vulnerable to noisy ECG curves and to variations in automatic trigging. That this approach is less than perfect can be seen in figs 1 and 3.
Limitations of the study
The present study is a feasibility study only, with limited numbers, but allowing a more detailed discussion of technical detail and individual findings. The prevalence of artefacts in a normal population, as well as the identification of those in patients, and the resulting diagnostic accuracy should be addressed in larger studies. It is still evident that interpretation of colour images is experience dependent, and influences reproducibility of diagnosis. Small numbers may also explain lack of significant differences shown in table 1, but findings that need larger numbers for significance have limited clinical value.
In this study, care was taken not to exclude artefacts from the quantitative analysis, to show the impact of those. As the study has demonstrated, colour imaging assists in identifying the artefacts, and the impact of identifying and eliminating artefacts before quantitative analysis has to be studied prospectively.
Post-systolic shortening, or thickening in the transverse direction  is generally defined as shortening after closure of the aortic annulus. In quantitative analysis, especially in strain, the exact location of end systole is important, to identify and measure peak value. In colour imaging, especially in 3D reconstruction, this is not equally important. The post-systolic shortening has a definite duration, extending into the early filling period, where elongation in other areas is visible. The exact location of the time of peak is not important.
Possible applications and further development
There are two main areas of application. In acute infarction, infarct area is important in prognosis. In addition, identification of an area at risk, may be important in clinical decisions, although the utility of this need to be established.
The other main area is in stress echocardiography. The experienced stress echo cardiographer does a quick visual assessment of the whole cine loop, before scoring in individual segments. This results in quick a decision reached of whether dyssynergy is present or not, and whether segmental scoring is necessary. Colour parametric imaging may to some degree duplicate this by making it possible to evaluate the homogeneity of colour, in one glance, and further to identify the segments where measurement is necessary. Colour display may also facilitate the learning of stress echo interpretation.
Processing tissue Doppler data from standard apical planes to a four-dimensional data set is quick, and feasible in all patients in this study. The tissue Doppler data can then be quickly post processed to any parameter, to be displayed in bull's eye, M-mode array or 3D surface; each display mode shows different information. Bull's eye displays the whole ventricle, while 3D surface reconstruction gives an approximate representation of true area. Data can also be re extracted from any point as waveforms for quantitative analysis. Parametric display allows quicker identification of both artefacts and dyssynergy than segmental quantitative analysis. Strain rate is best suited to parametric display. The display shows the area of post-systolic shortening in acute infarction exceeds the area of systolic dyssynergy, and may identify an area at risk. Further technical refinements are possible.
The study was in part supported by a grant from the Norwegian research council.
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