Relations between systolic
motion and deformation measurements
Apex to base differences
As the apex is stationary, while the base moves, the displacement and
velocity has to increase from the apex to base as shown below.
M-mode
lines from an
M-mode along the septum of a normal individual. These lines show
regional motion. It is evident that there is most motion in the base,
least in the apex. Thus, the lines converge in systole, diverge in
diastole, showing differential motion, a motion gradient that is equal
to the deformation (strain). This difference in displacement from base
to apex is also evident in the displacement image shown above.
 |
 |
| As the
apex is stationary, while the base moves toward the apex in systole,
away from the apex in diastole, the ventricle
has to show differential motion, between zero at the apex and
maximum at the base. Longitudinal strain will be negative (shortening)
during systole and positive (lengthening) during diastole (if
calculated from end systole). |
AS motion decreases from apex to base,
velocities has to as well. Thus, there is a velocity gradient from apex
to base, which equals deformation rate.
|
The systolic motion of each myocardial segment from the apex to the
base is the result of the segment's own deformation, added to the
motion that is due to the shortening of all segments apical to it.
Thus, as the apical segments shortens, this segment will pull on the
midwall and basal segments ( this is passive motion - tethering), the
midwall segment also shortens, and pulls even more on the basal
segment, which is shortening as well.
As the apical parts of the ventricle pulls on the basal, the
displacement and velocity increases from apex to base (
25). This means that
some of the motion in the base is an effect of the apical contraction -
tethering. In fact, completely passive segments can show motion due to
tethering, but without deformation. (
4,
6,
7), as also demonstrated
above. This means that the velocity (and
displacement) are position dependent, if not
normalised
while strain rate (and strain) are much more position independent, if
the velocity gradient is evenly distributed.
This is illustrated below.
 |
| Velocity,
displacement, strain rate and strain from three
different points, apex, midwall and base, in the septum of a normal
person. These curves all represent the same data set. It is evident
that motion
(velocity and deformation) increases from apex to base, showing a
gradient, while deformation (strain rate and strain) is more constant,
in fact a direct measure of the motion gradient. Diastolic deformation is far more complex,
and is discussed below.
|

|

|
Motion (left) and deformation (right)
traces from the base, midwall and apex of the septum in the same heart
cycle. It is evident that there is highest motion in the base (yellow
traces), and least near the apex (red trace), and this is seen both in
velocity (top - actually both in systolic and diastolic velocity) and
displacement (bottom). The distance between the curves are a direct
visualization of strain rate and strain, but the curves are shown to
the left, showing no difference in systolic strain rate or strain
between the three levels.
|
Is there an
apex to base gradient in strain / strain rate as well?
If systolic displacement and velocity decreases evenly from base to
apex, the
systolic deformation (strain) and velocity gradient (strain rate) is
evenly
distributed throughout the myocardium. Some of the earliest studies
seem to indicate this (
10,
19), although later studies seem to
find differences with
lowest values
in the apex (
124). However, the
angle error is also greatest in the apex (
206).
In the
comparative
study between methods in HUNT
(153)
there was lower values in the apex, but only using the longitudinal
velocity gradient, and only when the ROI did not track the myocardial
motion through the heart cycle. Thus, it seems fairly reasonable to
conclude that this finding is artificial.
With 2D strain, some authors have
found a ereverse gradient of systolic strain as
well, highest in the apex, lower in the base (207).
However, in that application, measurements are curvature
dependent, the apparent curvature being
highest in the apex and lowest in
the base, and the discrepancy between ROI width and myocardial
thickness being greatest. In addition, the strain values The HUNT study
(153)
found no such gradient with the combined
speckle tracking -TDI method, nor in the subset of 50 analysed for comparison of the methods.
| |
Basal
|
Mid
ventricular
|
Apical
|
Strain rate (s-1)
|
-0.99 (0.27)
|
-1.05 (0.26)
|
-1.04 (0.26)
|
Strain (%)
|
-16.2 (4.3)
|
-17.3 (3.6)
|
-16.4 (4.3)
|
Results from the HUNT study (153)
with normal values based on 1266 healthy individuals. Values are mean
values (SD in parentheses). Differences between walls are small,
and may be due to tracking or angular problems. No systematic
gradient from apex to base was found.
In addition, in the comparative study,
there was no gradient using the 2D strain application, in this case
care
was taken to align ROI shapes as much as possible.
MR studies have also found various results. Although some MR studies
have found a gradient, Bogaert and Rademakers (171)
found lowest longitudinal strain in the midwall segments, higher in
both base and
apex, but no systematic gradient from base to apex. MR tagging may have
some processing issues also, which may account for some of the findings
when curvature and angle varies long the wall from base to apex. Thus, the presence
of a base to apex gradient in deformation parameters has so far not
been established.
Differences
between walls
Although Höglund did not find any difference in systolic mitral
annular displacement between different walls (
30),
other authors have found such differences, with lateral displacement
higher than the septal (
167).
In the large HUNT study, the same differences were found in systolic
annular velocities (
165),
with differences between septum and lateral wall was of the order of
10%, but not in deformation parameters (
153),
where the same difference was on the order of 4% in strain rate and
only 1% (relative) in strain.
|
Anteroseptal
|
Anterior
|
(Antero-)lateral
|
Inferolateral
|
Inferior
|
(Infero-)septal
|
PwTDI S'
(cm/s)
|
|
8.3 (1.9) |
8.8 (1.8)
|
|
8.6 (1.4)
|
8.0 (1.2)
|
cTDI S' (cm/s)
|
|
6.5 (1.4)
|
7.0 (1.8)
|
|
6.9 (1.4)
|
6.3 (1.2)
|
SR (s-1)
|
-0.99 (0.27) |
-1.02 (0.28)
|
-1.05 (0.28)
|
-1.07 (0.27)
|
-1.03 (0.26)
|
-1.01 (0.25)
|
Strain (%)
|
-16.0 (4.1) |
-16.8 (4.3)
|
-16.6 (4.1)
|
-16.5 (4.1)
|
-17.0 (4.0)
|
-16.8 (4.0)
|
Results from the HUNT study (153,
165)
with normal values based on 1266 healthy individuals. Values are mean
values (SD in parentheses). Velocities are taken from the four
points on the mitral annulus in four chamber and two chamber views,
while deformation parameters are measured in 16 segments, and averaged
per wall. The differences between walls are seen to be smaller in
deformation parameters than in motion parameters, although still
significant due to the large numbers.
This is illustrated below.

|

|
Top: Pulsed wave recordings from the mitral
ring,
peak systolic velocity can be seen to be highest in the lateral
wall. Below;
the same can be seen in the M-mode recordings of the left
ventricle, lateral systolic annular displacement is seen to be higher
than in
the septum.
|
Top: Colour tissue Doppler recordings from
the same subject. Mark how colour Doppler recordings are analogous to,
but slightly different from the pulsed wave recordings to the
left. This is discussed more in detail in the ultrasound
section. The difference is also evident from the normal values of the HUNT study.
Below: Motion of
the mitral ring, can be shown by integration of the velocity, and both
peak systolic velocity
(top) and displacement (bottom) can be seen to be higher in the lateral
wall than in the septum.
|
 |
 |
| Deformation of the walls, both
peak systolic strain rate (top) and strain (bottom) can be seen to be
equal in the two walls (the small peak in the strain of the septum is
post systolic, and in addition only amounts to 1% absolute or 5%
relative). Thus the higher motion of the lateral wall is not
reflected to the same degree in deformation. |
The shape of the heart. As can be seen in
the top image, and is illustrated in the bottom, the curved lateral
wall is longer than the septum. Thus, strain rate (velocity
difference per length) and
Strain (shortening per length) is more similar between the walls than
just the
total shortening or velocity of the wall. |
This reflects the difference between motion and deformation on a
fundamental level, that
deformation
in the heart is motion normalized for heart size.
This is important both in
evaluating regional differences as well as global function, and is one
of the main advantages in using deformation imaging.
Tethering
The point of tethering it that a passive segment is tethered to an
active segment, and thus is being pulled along by the active segment,
without intrinsic activity in the passive segment. This means that a
passive segment may show motion, but without intrinsic deformation, and
the deformation imaging will discern. This is evident both in systole
and diastole.
tethering effects may show diverse results. It has three
important consequences:
- Infarcted segments may be totally akinetic, but still being
pulled along by active segments, showing motion without deformation.
This is usually evident in the inferior wall. A perfect example of a
totally passive, tethered segment moving close to normally, can be seen
below, and in more detail here. It may
also be pertinent to the basal part of the right ventricle. In both
cases, the annular motion may be near to normal due to hyperkinesia in
the neighboring segment, as this segment is offloaded as explained here.

|
|

|
The basal and midwall segment is infarcted,
and is akinetic and being pulled along by the active apical segment.
The whole inferior wall seems stiff.
|
Displacement
curves shows all segments to move similarly, thus there is little
differential motion, and at least below the apical point , little
deformation.
|
Strain curves, however, show that the
findings are more differentiated, showing akinesia basally (yellow),
hypokinesia in the middle (cyan) and hyperkinesia in the apex (red).
|
Thus; in this case, the passive segment is tethered,
showing motion and masking the pathology to some degree. Deformation
imaging will show this.
- If there is pathological contraction at some time in the heart
cycle (e.g. post systolic
shortening), the shortening of a pathological segment may impart motion
to a whole wall.

|

|
Velocity images showing motion towards the
apex in red, away from apex in blue. Left, systolic 3D
reconstructed image, showing normal motion in the septum and inferior
wall, and paradoxical motion in the inferolateral, lateral and anterior
wall. Right, om top are bull's eye from systole, showing the same, as
well as early diastole showing inverse motion during the e-phase, i. e
motion of the whole wall towards the apex in diastole. Apparently, the
whole anterolateral half of the ventricle is ischemic .
|
Strain rate images from the same recording,
left systole, right early diastole, showing that the ischemia is due to
a smaller ischemic area in the
inferolateral, lateral and anterior apex, where there is stretching
during systole (blue). This stretching, results in the midwall
and basal
segments moving away from the apex, despite contracting normally. In
early diastole there is recoil in the ischemic area (yellow), resulting
in
anterior diastolic motion in the whole of the wall. In this case,
the ischemia is obviously limited to a part of the apex, the rest of
the motion abnormalities being due to tethering.
|
In this case, the normal segments in the midwall and
base of the affected wall has abnormal
motion due to being tethered
to
the pathological segments in the apex. Another,
similar example of this in ischemia, can be seen below. Thus, it may mistakenly be taken ass asynchrony
between walls. Deformation
imaging shows the true location
and extent of the pathology.
- In phases where parts of the myocardium is active, other passive,
due to differences in timing, the tethering of passive to active
segments may make the whole myocardium move throughout the whole phase,
even if each segment is active only part of the time. This is evident
in diastole, where elongation occurs at different times in the
different levels of the myocardium.

|

|
(Motion (velocity), The diastolic phases of
early and late relaxation are seen as being simultaneous from base to
apex. Protodiastolic downward motion can be seen befor AVC (aortic
valve closure) in the tow basal segments.
|
Deformation (strain rate) shows both early
and late relaxation to be biphasic, and in addition the peaks are not
simultaneous in the different levels of the myocardium. Protodioastolic
elongation can be seen to be present in the midwall segment only, the
protodiastolic motion of the basal segment being a tethering effect.
|
This is explained in more details here.
The tethering
effects is the cause why
motion imaging
mainly shows global function, and deformation imaging shows regional
function. This has also been shown in a
clinical
study (
40).
Tehtering, however, being physiological phenomena, must not be confused
by the effects of
spatial smoothing
in
tissue Doppler or
2D strain. This is artefacts, being
due to shortcomings (or actually,; attempts to overcome shortcomings)of
the
methods themselves and
ultrasound in general.
Events of the heart cycle
The Wiggers cycle: Heart cycle in terms of
pressure changes
|
Volume and flow

|
| Classical Wiggers cycle, where events
during the heart cycle is related to pressure changes in atrium and
ventricle. The flow is a direct result of the pressure differences, and
thus the volume changes are the result of flow. It is evident that
pressure decline (relaxation) starts long before end ejection when
comparing with the image to the left. |
Top,
Ventricular volume through one heart cycle, with the different phases
demarcated. Below, composite Doppler flow velocity curve showing both
LVOT outflow and mitral inflow to the left ventricle. If the orifice
remains constant, the flow velocity will be similar to the flow rate
curve. Thus, the flow velocity curve is an approximation to flow rate,
and hence, similar to the temporal derivative of the volume curve, or,
conversely, the volume changes are the integrated flow rate. The
isovolumic phases are exaggerated.
|
Displacement and velocity

|
Strain and strain rate

|
| Top, mitral annular displacement curve,
being the curve showing the longitudinal shortening of the left
ventricle. Below, the tissue velocity curve, which is the temporal
derivative of the displacement curve. Comparing to the volume/flow
curve, it is evident that there is more complex motions, especially n
elation to the isovolumic phases, than is evident from the mere volume
diagram to the left. |
Top, strain curve from mid septum, showing
the deformation, below the strain rate (temporal derivative). The
curves seem to be very similar to inverted motion and velocity curves,
however, deformation will show more regional detail as discussed below.
Remark also how the strain curve is similar to the volume curve,
showing the same pattern, while the strain rate (temporal derivative of
strain) is similar to the flow curve (temporal derivative of volume).
|
It has been established that the
longitudinal shortening of the left ventricle, and thus the
longitudinal
measures is closest related to the stroke volume and EF, i.e. to the
total left ventricular volume change (
13,
30
- 35,
56,
59,
60,
64 -
67,
116).
Thus, the longitudinal strain is the most important measure, and it is
also closely related to the wall thickening and thus internal
shortening as discussed
above.
Systolic events
The difference in shape between strain
rate and velocity curves
It is obvious that strain rate and velocity curves are different. Apart
from being inverted, which is due to the
subtraction
algorithm, the systolic strain rate curves are much more rounded,
with a later peak, while velocity curves show a sharper and earlier
peak. If strain rate is equal to
normalised velocity, why is the shape different as illustrated below?

|

|
Left velocity curves. It
can be
seen that the two velocity curves have an early maximum, showing that
the myocardial acceleration occurs early, and is an early event. Peak
systolic velocity is seen at about 100 ms into the heart cycle,
starting with ORS. After this, there is a period of nearly
constant velocity difference, before the velocity difference decreases
again. Right,
the strain rate curve from the segment between the two ROIs in the
left picture. It can be seen that peak strain rate is a later
event, about 200 ms after start of QRS. The strain rate increases
after peak velocity, during the period of near constant velocity
difference. This is due to the fact that
the velocity difference is
normalised for the instantaneous distance that is decreasing during
systole, i.e. Eulerian
strain rate.
This is due to the difference between Lagrangian and Eulerian strain
rate, which is explained in detail
here.
As we use
Lagrangian strain, this is
displacement normalised for end
diastolic length. This is the original definition, and the one used to
describe myocardial deformation by Mirsky and Parmley (
12),
and thus has set the standard. However, it has become customary to use
Eulerian
strain rate , which is a normalization for instantaneous length. The
reason for this is that this is equal to the velocity gradient that was
the original method used to calculate strain rate (
4,
14)
and
is what
is used.directly on the scanner.
This means
that the distance between the points of velocity measurement decreases
during the whole systole. Thus, after the peak velocity, while the
phase where velocities are relatively constant, strain rate
will continue to increase as the strain length decreases. Thus, peak
strain rate is a later
event than peak velocity, which means that it
may be more load
dependent than peak systolic velocity. In addition, the Eulerian strain
is slightly higher in value than Lagrangian strain as discussed
here,
as the systolic
deformation is normalised to a decreasing in stead of a constant
length. This, however, means that peak Eulerian strain is the highest
contraction velocity (not rate, as contraction rate should be measured
in terms of tension). Peak velocity (or Lagrangian strain) is probably
closest to the time of maximum dP/dt, while peak Eulerian strain rate
may be closest to the time of peak pressure.
The time of peak strain rate has not proven useful, however, as the
method is so noisy that timing is more influenced by noise spikes than
the true time point of maximum strain rate.
Lagrangian
strain (
) and Eulerian strain (
) can be
interchanged by fairly simple conversion formulas:
and
.
The convention is that the
strain is given as Lagrangian strain. Integration of Eulerian strain
rate yields Eulerian strain, but this is converted directly to
Lagrangian strain by the formula above, so the curves and values seen
on the workstation are Lagrangian values. Conversion between strain is
thus useful, and used all the time, even manually. The main point of
interest is end
systolic strain or peak systolic strain which will be simultaneous by
both measurements. (Of course, to obtain a full Lagrangian strain curve
from Eulerian strain, the correction has to be applied in each frame,
which is a little more computationally expensive, and thus needs to be
automated in the analysis program. However, peak Lagrangian strain rate
will be at the
time of the biggest velocity difference, while peak Eulerian strain
rate will be later. Thus, peak strain rate is not simultaneous by the
two variables, and peak strain rate with one method will not convert
into peak strain rate of the other. Thus conversion of strain rates is
practically useless for peak
values. However, the conversion has to be applied to each frame if Lagrangian
strain is obtained by speckle tracking, and then derived to achieve
strain rate. This is also done automatically in the analysis programs.
When is aortic valve closure in relation
to the events seen in echo?
The timing of end systole is crucial to defining end systolic strain,
especially in the cases of post systolic shortening. End systole is
often defined as
end
ejection, as defined by the aortic valve closure (AVC), as shown
in the diagrams
above. The end
ejection is easily seen in Doppler flow recordings from the LVOT, by
the aortic valve click, as described
here.
Parasternal recordings of the aortic valve can also identify the AVC,
but due to the longitudinal motion of the heart, the aortic valve often
moves out of the M-mode line in end systole. However, at the present
stage of technology,
Doppler flow recordings must still be taken separately from B-mode
recordings, with or without tissue Doppler data. By transferring the
AVC from a Doppler flow recording the heart rate variability may lead
to errors in the estimate of the AVC, as the ejection time is
proportional to the total cycle length (RR - interval) (
29). The ECG has a low precision
in timing end systole, and regression equations based on heart cycle
length has limited validity as the relation between RR interval and
ejection time is not linear, at least not over the full range of heart
rates (
29). By interfacing a
phonocardiograph with the scanner, the timing of valve
closures can be done in all recordings. However, low level noise may
lead to small errors in detecting the earliest part of the first heart
sound, and so the phono should be calibrated by Doppler.

|

|

|
Apical recording of Doppler flow of the
LVOT. At end ejection, the valve click can easily be seen as the short
spike. This is coincident with the start of the phonocardiographic
first heart sound as seen by the phonocardiogram. However,
in the last heart cycle, there can be seen a small oscillation
earlier in the others, a small noise spike (red arrow). Thus the
Doppler is the gold standard, and the phono has to be calibrated.
|
Parasternal
long axis of the aortic valve, Due to the longitudinal motion of the
base of the heart, the valve has moved out of the M-mode line at end
ejection, and the AVC cannot be seen.
|
But this recording was done with tissue
Doppler superposed, and turning on the colour reveals the valve click
as a vertical blue line (marked by the yellow arrow). The visibility in
tissue Doppler is due to the broader beams and different filter setting
of tissue Doppler compared to the B-mode.
|
The tissue velocity traces shows a small and short negative spike at
end ejection. This was early assumed to be the isovolumic relaxation
resulting in a shape change of the left ventricle. The AVC was thus
assumed to be at the start of this spike by various authors. This
negative event can also be seen in colour M-modes of tissue Doppler,
both in the mitral ring and the mitral leaflet. The negative spike will
then correspond to a narrow blue (negative colour) band, and the AVC
was assumed to be at the start of this band. This has even been
published as a method for determining the timing of AVC in tissue
Doppler images. This is illustrated below.

|

|

|
Short, negative velocity spike at end
ejection. This ha erroneously been assumed to be isovolumic relaxation,
and
hence, AVC at the start of the spike.
|
The negative spike corresponds to the
vertical narrow blue band (blue = negative velocity) and perpetuating
the mistake, the AVC would be at the start of this blue band as marked
by the black arrow.
|
Locating the AVC by this assumption, the
method of tracing an M-mode across the anterior mitral leaflet has been
published.
|
However, as one knows that there is relaxation during the last period
of ejection, it is conceivable that there is a small elongation at end
ejection that stops abruptly with the AVC, which is a sharp, mechanical
event. This can be seen in both parasternal m-modes of the septum, as
well as longitudinal M-modes of the mitral ring and mitral ring
displacement traces. In addition, using high frame rate imaging of the
septum together with a high fidelity phonocardiograph, we could se this
elongation before the AVC in an observational study (
16).

|

|
 |
Well known finding of a systolic "notch" in
the septum in systole. This corresponds to a slight thinning of the
septum with an abrupt stop.
|
Displacement curve of the mitral ring. (The
same can be seen in M-mode). It can be seen that there is a short
motion away from the probe, corresponding to the negative velocity
spike at end ejection. The motion stops abruptly, and there is a slight
"bounce" before mitral opening leads to another downward motion.
|
Colour M-mode from the
septum of a normal
subject. It is evident that there is an elongation in mid septum,
resulting in initial
negative velocities in mid and
basal septum before
closure of the aortic valve. Notice also how the
initial elongation of the mid septum occurs before the closure of the
aortic valve, i.e. the initial negative velocities in the basal and mid
septum are protodiastolic.
|
Using first phono that was calibrated by Doppler, we were able to show
that the observation by strain rate imaging was actually true. AVC was
in fact at the end of the negative spike, where velocities crossed back
from negative to positive, i. e. corresponding to the "notch" in the
mitral ring motion (168). Although
for practical purposes, the automated algorithm identifies the point of
maximum acceleration, which is very close. Later we used a scanner that was modified to acquire B-mode and tissue
Doppler alternating in an 1:1 pattern, and in narrow sector of the
septum giving a frame rate of close to 150, imaging both the base of
the
septum and the aortic valve at the same time in 5-chamber and
long axis views. Here, the actual closure of the AVC could
be identified with a temporal resolution of about 7 ms. The study
confirmed the previous findings (169),
and, repeating the study in infarction patients and in high frame rate
during stress echo, showed the finding to hold true (170) also outside the normal range.
Thus, The initial negative spike is a protodiastolic event, the
continuation of relaxation into the phase after the final shortening,
as shown below. Thus, it is not a
measure of isovolumic relaxation. However, this still means that
the AVC can be
identifies as a mechanical event without recourse to the flow curves or
direct visualization of the aortic valve.

Correct positioning of the AVC
by tissue Doppler of the septal mitral ring is where the protodiastolic
negative velocity spike crosses zero, and becomes positive.
The AVC should preferably be located from the basal septal traces, as
the closure of the aortic valve is a mechanical event that propagates
through the myocardium, and thus will be slightly later with increasing
distance from the aortic valve (towards apex and in the lateral wall),
as shown by high frame rate TDI (
172).

|

|

|

|
Placing the AVC event marker, shows the
protodiastolic negative velocities to be present in the basal and
midwall segments (yellow and cyan curve), but not in the apex (red
curve). Converting
the dataset to a curved M-mode, the spike corresponds to the narrow blue
band, and the zero crossing to the shift from blue to red.
|
Keeping the event marker, but converting to
displacement, wee see the "notch" in the basal (yellow) curve, and the
AVC is the bottom of the notch where there is an abrupt change from
downward to upward motion, thus the change from negative to positive
velocities.
|
 |
 |
 |
| Keeping the event marker in place, but
converting to strain rate and strain. Now it can be seen that there is
an elongation only in the midwall (cyan curve). The finding of negative
velocities in the base as well, is due to tethering, and shows how
deformation imaging has a better spatial resolution in separating
events in space. If AVC should be placed by strain rate
traces, it can only be located from the M-mode or the midwall race,
just
after the initial elongation, but strain rate traces shows a generally
complex pattern and are little suited to location of AVC. M-mode is
far better. In strain curves, however, AVC can again be seen as a
"notch" (as in displacement),
most evident in the midwall (cyan) trace. |
Thus: three things are evident:
- Looking at elongation, there is an initial elongation in the
midwall before the AVC. This has some bearings on the mechanism for the
aortic valve closing as shown in the illustration below.
- The AVC can be located in most traces (provided a sufficiently high frame rate),
without transferring data from a Doppler or phono recording, preferably
in the septum, and most easily in the basal segments of motion traces
or the midwall segments of the deformation traces.
- The midwall elongation is not "strain during IVR", and it is
mainly present in midwall, and the amplitude is position dependent.

Proposed mechanism for the
aortic closure. During ejection the ventricle can be seen to shorten,
and there is ejection (arrow), keeping the cusps open. Ejection
is decreasing towards the end of the ejection period, as shown by
the decreasing length of the arrow. At end ejection, there is no flow,
and the relaxation
that started during ejection as a reduction in tension, leads to a
slight elongation. The aortic cusps then are closed due to
the action of the now stationary blood column, similar to what happens
if a scoop is put into the water (opening forward) from
a boat that is moving forward. The aortic valve stops when the cusps
close, there being no further room for backward motion. This leads to
an abrupt stop in the motion of the base of the heart, and a small
"bounce", which is what's seen in the motion traces above. (The
"bounce" is not depicted in the animation.)
This model of early relaxation was later confirmed by a
combined experimental and theoretical analysis (
173), although the interaction with
the blood column was not specified, and the load dependency of early
diastolic tissue velocity was taken to mean that the load (filling
pressure) was part of the mechanism for ventricular elongation
(enlargement), although this is doubtful, considering that the pressure
in the ventricle actually drops during early diastole as discussed
below.
Diastolic events
When is mitral valve opening?
The opening of the valves is a passive event, where the valves follow
the blood flow, with the same motion and velocity. Thus the valve
opening is the start of flow through the valve. As with AVC, due to
heart rate variability, it would be advantageous if the mitral valve
opening (MVO) could be identified in the tissue Doppler traces, instead
of being transferred from other cycles.
The logical candidate would be the moment the mitral ring starts to
move away from the apex, after the "bounce" following the AVC. This
would hypothetically be the time point where the ventricle starts the
volume increase, seen as elongation by the mitral ring. But as this is
not an abrupt mechanical event, but rater a gradual transition (an
upwards convex curve in the displacement traces), this is not as easily
delineated. However, this will correspond to a shift from positive to
negative velocity in the velocity traces ans a red to blue shift in the
colour traces, as seen below.

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Mitral opening possibly corresponding to
the start of elongation after the MVC. This means again a shift from
positive to negative velocity.
|
Mitral valve opening identified by the
shift from positive to negative velocity. It is the second zero
crossing after the protodiastolic dip. Event marker may placed by velocity, and
then carried over to the displacement traces, as the zero crossing
point may be easier to identify than the transition in the curvature in
the
displacement trace. However, as seen from the M-mode, the
transition from positive to negative (red to blue) is not evident at
all levels.
|
But the exact time point is not as
well
defined, and it may correspond to another event, such as the point of
maximum negative acceleration. So far, the positive to negative shift
in velocities is an approximation.
What about the mitral valve itself?
As the mitral valve is visible in all apical views, it should be
possible to identify the mitral valve opening directly. However, as the
base of the heart moves slightly towards the apex after AVC, the mitral
valve motion follows the mitral ring. When the mitral valve opens, flow
starts and the ventricle expands (elongates) corresponding to the
downward shift in displacement of the mitral ring. However, the mitral
valve opens, meaning motion of the leaflets into the ventricle,
continuing the motion towards the
apex. Thus the leaflets do not have the shift from positive to
negative velocities. Some authors have described this
anyway, but this is due to the fact that the
lateral resolution in tissue
Doppler is very low due to the low line density, in order to achieve a
high frame rate, meaning
that an M-mode line placed across the mitral leaflet close to the ring
actually will be ring velocities as discussed in the
pitfalls
section. In addition, the base of the mitral leaflets will tend to
follow the ring motion more than the tips. Also, the opening of the
mitral valve is gradual, starting at the tips, and moving outwards
towards the ring.
The
aliasing velocities of the ring are even later, as this marks
the time when the leaflet movement (moving with the same velocity as
the flow, as discussed
here)
reaches the aliasing velocity of the tissue Doppler, being dependent
on the PRF and depth. This is also earliest at the mitral leaflet
tips.

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 |
 |

|
Motion of the mitral ring, mitral leaflet
and mitral tip. Bottom; zoomed to the time period of interest.
The mitral ring (yellow curve) can be seen to "bounce" after AVC.
The middle part of the mitral leaflet (green curve) can be
seen to be partially following the mitral ring, and motion towards the
apex starts before MVO. The mitral tip does not follow the
"bounce" of the ring (probably due to the inertia of the blood in the
cavity, and in fact there is a slight billowing of the valve in the
middle, evident by the continuing motion away from the apex). The
motion towards the apex starts simultaneously with the
motion of the ring away from the apex. However, here, the "notch"
marking AVC is also absent, but the change from negative to positive
velocity should be evident. Both mitral valve traces can be seen
to deflect sharply downwards at a later time point (white markers) ,
this is due to aliasing of the tissue velocity when the velocities
reaches the Nykvist limit.
|
Velocity
traces of the same points as seen to the left. The mitral tip can be
seen to cross from negative to positive velocities at the same time as
the mitral ring crosses from positive to negative. The points of
aliasing can be seen at the abrupt downward stroke in the traces from
the mitral leaflets (White markers), which is earlier at the tip than
in the middle of the leaflet. Only in the mitral ring can AVC be seen
with certainty.
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M-modes from the mitral ring (left), middle
mitral leaflet (in terms of distance from ring to tip - middle image)
and mitral tip (right image), corresponding to
the traces above. The traces show only systole and early
diastole, as above. As in the traces there is shift from negative
(blue) to positive (red) at AVC and from positive to negative at the
event taken to be closest to MVO. In the middle mitral leaflet,
however, this is less well defined, only in the lowest part can these
event being identified. At the mitral tips, the transition marking AVC
is absent, (as above), while the second transition from blue to red is
visible in a short part of the M-mode.
|
From the images above, it seems that the mitral tips billow away from
the apex during the IVR (as expected), and the starts to move towards
the apex at MVO. Thus registrations from the mitral tip might be
feasible, but this should be done with high frame rate. However, high
framerate leads to less lateral resolution, thus mixing the lateral
parts and the tips. Narrow sector in the middle of the mitral tips
might be used, but if the time point then needs to be transferred to
another full sector cycle, the MVO time could just as easily be taken
from pwDoppler of mitral flow. Thus no reference for the ground truth
can be established within one and the same heart cycle with any
reliability.
What about strain rate and strain
curves?
As shown
above and discussed
in detail
below, the strain rate
curves show a very complex pattern, and is unsuited for locating events
in this part of the heart cycle. Also the strain curves shows different
patterns in different levels of the myocardium. Thus, the deflection
points can be seen to be located differently in the different levels of
the wall. In addition, the presence of post systolic shortening,
especially in pathology, but also in normal ventricles, will result in
the shortening will last longer in the strain rate then the
upward movement in the displacement.
Mitral valve opening should thus be
identified in velocity images and transferred to deformation images.

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Zoomed images of velocity (left) and
displacement (right), showing that there is a
peak apical mitral ring displacement at the event taken to
correspond to MVO. This is equivalent with the second
zero crossing of the velocity trace after protodiastolic dip.
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 |
Peak negative strain occurs later in base
and
midwall. AVC can be seen best by strain curves in the midwall segment.
No definite deflection can be seen to correspond to the event assumed
to be MVO transferred from the Velocity/displacement traces.
|
Looking at the velocity and
displacement traces, even with the addition
of the protodiastolic motion event, the diastole looks fairly
straightforward, after AVC, the three fundamental phases known from
Doppler flow can be seen: Early filling phase (E), seen as the first
negative phase (e') after AVC, diastasis with little or no motion, and
the atrial systole (A) seen as the second negative velocity spike (a').
The atrial displacement of the ring may be described as the atrium
pulling the ring away from the apex, and in addition the added volume
pushed into the ventricle by atrial (esp. auricular) contraction
pushing the atrioventricular plane. The relative contribution of the
two mechanisms is uncertain.
Taken from the mitral ring, diastolic ventricular displacement and
velocity show the left ventricular
diastolic
global
function.
 |
 |
|
Velocity and
displacement in the base of the septum, showing systolic motion toward
the apex, protodiastolic motion away, and the the two basis diastolic
phases, early (e') and late (a') motion way from the apex, separated by
diastasis.
|
However, using strain and strain rate, the
diastole can be seen to be far more complex, showing a sequence of
events that are different, and with different timing in the different
segments. Thus is seen due to the better spatial resolution, as
deformation imaging eliminates the effects of the tethering of the base
to the more apical parts. In addition, these events interact, to result
in the simpler pattern seen in motion traces, and the main finding is
that there are more than one peak in each of the two phases of E and A,
and also, the peaks are not simulatneous in all parts of the ventricle.
The double peak of Una's tits
(yes, that is the actual, official name of this mountain) may be a
reminder of the double peaks of diastolic strain rate.
|

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| In strain and
strain rate, the pattern can be seen to be much more complex in these
tracings from the base alone. There are at least four positive spikes
(elongation) during diastole, this is reflected by much more "steps"
towards zero in the strain curves. As strain rate is
fairly susceptible to
noise, this might have been interpreted as noise (as is the small
negative spikes between) , but integrating to strain eliminates the
random noise, and shows what is real. |
The
protodiastolic phase cannot be seen in the traces from the base, only
in the mid wall in the M-mode. Also the M-mode reveals that some
of the diastolic phases has the characteristics of elongation waves
between base and apex. The differences between base, midwall and
apex can be seen clearly in the traces diagram to the left, showing
that not only are there more elongation phases, but they don't even
coincide in the different levels of the heart.
|
After the protodiastolic elongation,
there is a diastolic elongation
that is most evident in the apex. This occurs before the opening of the
mitral valve. The opening of the mitral valve signals the main
elongation (and wall thinning) that starts at the base (
19).
Tissue
M-mode from the septum, showing the dip of the AVC event, and the time
delay from base to apex of the initiation of downward motioning the
filling phase.
The relation of this propagation to diastolic function is discussed
below.
But as this wave has propagated to the apex, it can be seen to return
to the base. And the elongation during atrial systole can be seen to
behave similarly with a wave toward the apex, and back to the base, but
with a higher propagation velocity that the early relaxation, which
makes sense considering that during early filling the ventricular
myocardium is in a relaxing, and in the atrial systole, the ventricular
myocardium is relaxed.

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 |
Events of the heart cycle seen by strain
rate. Both the curved M-mode and the traces shows the separation into:
- Midwall protodiastolic
lengthening
- Apical isovolumic
lengthening
- Early filling propagating
from base to apex and back
- Late filling propagating
from base to apex and back.
|
Proposed explanation of the return wave of
the two filling phases, in reality being a crossing over from the
opposite wall. The protodiastolic lengthening is less evident in the
lateral wall, but this is due to a drop out in the wall.
|
The finding of a complex pattern in diastole, shows that
no single parameter can be used as a
criterion for diastolic function. Regional early strain rate
might be taken as an indication of regional diastolic function, but
only if care is taken to identify the elongation spike, and avoid the
return wave. And as the traces above show, there are differences in
both the amplitude and timing of early diastolic strain rate, the
implication being that there is no meaningful way of averaging the
values into a more global function measure. The e', being the resultant
velocity of the mitral plane, however, is a truly global measure, being
the summation of all local measurements and taking the time differences
into account, as well as being less pressure dependent, is a more
robust
measure of diastolic function as discussed
below.
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 |
This
picture is too busy to be really informative, it summarizes the
information given above, and is mainly a way of showing that there
is no meaningful way of averaging the diastolic peak values into a more
global
function measure, as the peak local diastolic strains are not
simultaneous.
|
For global diastolic function, diastolic
tissue velocity is still .the most important measure, as this is the
resultant global peak measure. This is not the average, but the
resultant of all the local (non-simultaneous) diastolic strains AND the
propagation along the wall.
|
On the other hand, the finding of the
filling phase as a propagation wave from base to apex shows another
measure of diastolic function, as well as a different relation between
tissue velocities and strain rate as shown
below:
Strain and strain rate in the atria
As
the outer contour of the heart is relatively constant, the
apex is stationary, and the atria is attached to the large veins, the
atrioventricular plane has to be the piston of a reciprocating pump as
discussed
above), expanding the atria while
the ventricle shortens and
shortening the atria while the ventricle expands. This is energetically
useful, as the work used to decrease the volume, in additon to
ejection, also moves the blood from the veins into the atria. If the
heart had worked by squeezing changing outer contour to a high degree,
the work would have been used to shift the rest of the thoracic
contents especially lungs inwards in each systole, work that would have
been waisted. Thus, most of the filling volume to the ventricles, is a
function of the AV-plane pumping.
Basically, the deformation of both chambers reflects the motion of the
atrioventricular plane. In systole, the ventricle shortens while the
atria expands. This is a function of ventricular contraction. In early
diastole there is elongation of the ventricles and shortening of the
atria, the active component of this is the ventricular relaxation. In
late diastole, there is further elongation of the ventricles and
shortening of the atria, but in this phase the active component is the
atrial contraction. However, deformation of both chambers are
reciprocating, both reflecting the atrioventricular function, and
for the elongation of the atria during ventricular systole is not
an
independent parameter, and is mainly due to the systolic function
(shortening) of the left ventricle.

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| Deformation
in the atria is reciprocally related to the deformation of the
ventricles. as both apex and the atrial roof are relatively
immobile, the ventricle shortens (yellow) while atria elongates
(cyan) during systole, ventricle elongates (cyan) and atria shortens
(yellow) during the diastolic phases. In diastasis, there is no
deformation, both are green. |
Curved M-mode going
through both ventricular and atrial septum shows the reciprocal colours
of the atria and ventricles. |
Strain in atrium
(yellow) and ventricle (cyan) are seen as almost mirror images of each
other. However, the absolute values in the atria are higher, as the
atrioventricular plane motion is a greater percentage of the smaller
atria. Strain rate curves are also
basically mirror images of each other. As deformation is active in one
chamber and passively transmitted to the other, the peak values may be
higher in the active chamber, and there will be a time delay of events
as waves propagate as shown in the ventricle during diastole. |
Thus, deformation during ventricular systole, reflects ventricular
contraction, early
diastolic deformation reflects ventricular deformation and A wave
reflects atrial contraction, irrespectively of which chamber the
measurements are done, although values may vary. as shown below.
Thus: The AV-plane motion is
determined by the systolic function of the ventricles. The atrial wall
being thin, there is little resistance to AV-plane motion, and does not
express distensibility of the atria. If one choses to measure strain in
the atria instead of the ventricles, the ventricular strain being
MAE/LV length, the atrial strain will be MAE/ atrial length. Thus the
atrial strain will reflect the relative change in LA length, and thus
the "reservoir function", but says nothing of the atrium itself.
In
this subject there is a ventricular systolic strain of 15%, while
atrial strain during ventricular systole is 38%. However, taking the
different lengths of the atrium and the ventricle, and calculating the
absolute change in length, it can be seen to be the same within the
limit of accuracy. This is simply the MAE, reflecting both
shortening of the ventricle and (longitudinal) expansion of the atrium.
Atrial strain during ventricular
systole
Reduced "reservoir
function" (atrial strain
during ventricular systole) in the atria in atrial dilation, is simply
a function of the increased atrial length, given normal ventricular
systolic function. And this will be
relative, the absolute increase is the same, given the same MAE. In
fact: the same MAE will result in the same
absolute longitudinal
expansion of the atria, and thus, much the same volume increase
irrespectively whether the atria is long or short.
The relative
expansion, however has different contents.
- Atrial strain is MAE divided by the atrial length. Thus, in
reduced systolic function, the MAE is reduced, and so is atrial strain.
- In atrial dilation, the atrial strain during ventricular systole
will be reduced even with normal MAE, as atrial length is increased.
Thus atrial strain during ventricular
systole is a composite of longitudinal ventricular function and atrial size. Left
ventricular filling pressure (LA pressure), is elevated in reduced LV
function, especially in heart failure. Thus, studies showing reduced
left
atrial strain with increased filling pressure may thus suffer from
confounding with LV function.
Also, the left atrial size per se, is
a sensitive index of chronic atrial pressure over time (
195), and thus, even in normal LV
function, the LA strain may correlate with LA pressure (and indeed may
be a function of LA pressure), but the parameters are not independent.
Thus, an independent contribution of LA strain to the echo evaluation
may be doubtful.
Being a composite parameter, it may be
more sentitive than sigle parameters, but not independent. Long
axis shortening is the most sensitive parameter of both function and
prognosis (36, 190, 191, 192, 193), far morte than EF. Thus, in
comparing the efficacy of different parameters, the longitudianl
shortening should be used as systolic measure, and included in multi
variate analyses in order to determine whether atrial strain during
ventricular systole actually gives independent information.
Furthermore:
- The longitudinal expansion of the
atria corresponds to a greater volume dependent on the width
(cross
sectional area/diameter). The atrial expansion times the cross
sectional area, would thus reflect a part of the reservoir volume in
the atria. However, the auricles would not be included, and the volume
function would not be complete.
The wall length should probably be related to a curved strain
length (as in 2D strain). However, because of its thin wall, strain
rate imaging in the atria
from the apical position is extra prone to artifacts due to
low lateral resolution at the
large depth, which may affect transverse tracking in the atria, and
result in under estimation of the wall shortening anyway.
Thanks to Mads Ersbøll of
Copenhagen, who pointed out the connection to LA pressures in studies,
leading to the discussion in this paragraph being extended,
incorporating the relation to atrial size.
Atrial strain during early diastole
Early filling phase is likewise related to ventricular
diastolic function, the mechanisms
being elastic recoil modulated by the rate of calcium removal from the
cytoplasm as discussed
below.
Thus, the amount of the systolic atrial strain being reversed in early
diastole is also a property of the ventricle, divided by the atrial
length.
Atrial strain during late diastole
Finally, the atrial contraction is a property of the atria.
It has been proposed that the main function of the atrial systole is to
pull the mitral ring back to the end diastolic point, thus pulling the
mitral ring over a volume of blood contributing to the end diastolic
volume(
13). However, this model
disregards the finding by Doppler flow that there is an active flow
component as well. And, as the pressure in the ventricles
increase during the atrial
systole, there is evidence for the vis a tergo mechanism being an
important component also of the motion of the mitral ring, i.e.
pressure being the driving force. Thus, the peak A is the volume flow
due to the contraction of the atrium, but modified by the rate of
pressure increase in the LV, being a function of the LV compliance.
However, the AV-plane motion and a', would be measures of atrial
function. And, as this is pumping volume driven the total atrial action
(including the pumping of the auricles) would be the driving force.
As e'/a' ratio decreases, the a increases, so this function is not
independent on LV function, but comes closer than other measures. And
is of little use where there is partial
fusion of
e and a. as the velocities then are combined partially of
ventricular relaxation.
Atrial strain and strain rate are simply displacement and velocity
normalised for
atrial length
as shown below.
The true atrial contractile
function is the length change during atrial systole. Changing hte start
of tracking to the start of the a-phase, shows atrial strain to be 13%.
Peak strain rate can be measured independent of the starting point for
tracking.
Thus: the atrial long axis function by deformation, basically reflects
the AV plane motion. The added value of using of atrial strain and
strain rate instead of annular displacement during atrial contraction
is uncertain. However a theoretical advantage is that it corrects for
atrial dilation, and thus mau be more sensitive, but this needs to be
estabolished in clinical studies.
Global systolic function
With the appearance of new methodology, a number of new methods for
measuring left ventricular global function has emerged. Older measures
has traditionally been measurements of the
cavity
function: Stroke
volume, ejection fraction (and the M-mode equivalent shortening
fraction). Newer methods include
longitudinal measures
of wall function, as annular displacement and velocity, as
well as mean strain/strain rate, either based on segmental
measurements, or a global averaging (as global strain form speckle
tracking
2D
strain). It should be
of general interest to comment on the relationship between the methods.
It is also important to realise that while strain and strain rate are
measures of shortening per length unit, the annular velocity and
displacement are also measures of the same, but in absolute values
(i.e. not normalised for ventricular length). However, all measures
that measure relations to changes, i.e. in paired experiments of load
alterations, the normalisation will cancel out, and displacement will
behave as strain, strain rate as velocity (more or less see
difference between Eulerian and Lagrangian strain
rate). Thus all experiments with systolic displacement and velocity
in relation to global changes, will pertain also to strain and strain
rate.
What does Strain and strain rate actually
measure?
It is important to realise that strain and strain rate measure
only deformation.
Deformation is definitely NOT contractility. In fact,
any imaging technique, by
any parameter measures
deformation
which is the visible contraction, not contractility. This is due to
more than one point:
- The greatest apart of the work is the pressure buildup, during
the isovolumic contraction, i.e. the pressure work,
which is mainly isometric. This cannot be measured by imaging
(including deformation imaging) at all.
- Deformation of the ventricle during ejection, i.e. the
contraction is load independent, the work being a function of both
deformation and prerssure.
- Basically, contractility is described in terms of the
volume-pressure relation, meaning that the increase in contraction due
to the Frank-Starling mechanism is excluded
So any imaging measurement
will measure
load dependent
contraction, unless some correction is done
to obtain a measure of load.
(As contractility in fact is the development of force, the most direct
measure should have been strain rate acceleration, acceleration being
directly related to force. However, as strain rate is a fairly noisy
method, derivation to strain acceleration have so far been shown to be
prohibitive because of noise. And still, it would only be the force
leading to deformation, not pressure build up.)
Contractility is the
ability to develop force independent on load, and is closely related to
the stress-strain
relationship. I.e. the shortening in relation to the load. (Thus, for
any given force, the deformation is load dependent, but as force may
change with both heart rate (force-frequency effect) and volume
(actually initial load; the Frank-Starling effect), the interaction may
be quite complex.
Contractility may be defined as end systolic (actually end ejection)
pressure volume relation, (even though this changes with inotropy, this
is slightly dubious as the myocytes actually are in a relaxing mode at
that time). An other measure is the peak dP/dt. This occurs during the
isovolumic contraction period, when there is no deformation. Thus this
is a simple measure of force development, but it's still dependent on
preload, which is a function of volume
(diameter) and end diastolic
pressure.
But, again, this the part of work related to pressure buildup, the
pressure work. Considerable energy is
used to build up the ventricular pressure from the low filling
pressures of the left atrium to the high ejection pressures of the
aorta in systole. This may be considered as
the
mainly isometric part of LV
work, as there is little deformation. The
strain and strain rate
are both indices of ejection work. In this phase, there is significant
reduction of volume, but far less change in pressure. So this may be
considered a more isotonic component of the work. And it is only this
part that
is measured by the systolic deformation indices. As described above,
there may be differences in the different fibres in the contribution to
the two parts of the work.
Thus, to lump everything together and calling it contractility, may be
meaningless, there may be a correlation, but the concept of
contractility, outside of isolated muscle experiments, is more
theoretical. Strain and
strain rate measures deformation, and deformation has to be interpreted
in terms of load, to infer contractility.
Basically, longitudinal strain and strain rate are methods to measure
regional deformation, the
basic algorithm subtracts the motion due to contraction of neighboring
segments (tethering effects). In principle, velocity and displacement
measures the effect of contraction of the whole ventricle apical to the
point of measurement. Thus, annular plane displacement and velocity
measures the
global
function of the left ventricle (
13).
This has been demonstrated in several studies, both for systolic
annular
displacement (
30 - 36)
and velocity (
37 -
40).
Thus, as deformation is a result of tension, or rather tension versus
load, strain does not measure function directly. But taking regional
function into the concept of load, deformation imaging can be used to
infer force, or
at least inequalities in force development, as shown
below. Thus, strain rate images shows
gradients of
relative
contractility, even if one does not measure absolute
contractility.
And that is the main point in
regional diagnosis.
Cavity measurements of systolic function
Ejection fraction
Based on Nuclear or X-ray contrast studies, the first measures was
measurements of cavity reduction in systole, i.e. the stroke volume.
While this may be the most important
result
of cardiac pumping, it confers little information about the state of
the heart itself. A dilated ventricle can maintain stroke volume, but
it is reduced in terms of the left ventricle volume, and may have a
severely reduced contractility. Thus stroke volume should be normalised
for end diastolic volume, to obtain Ejection fraction:
Ejection fraction is still the most
widely used measure of systolic left ventricular function today. This
is mainly due to the vast amount of prognostic information from earlier
studies, and the prognostic interventions that are geared to a cut off
point in EF. Thus it will remain in use for an foreseeable future. But
alas, interventional studies using echocardiography as secondary
outcome, persists in using only EF, instead of including newer measures
for direct comparison of the ability in predicting clinical outcome as
well as establishing cut off values for intervention.In
assessing EF, it should be emphasized, however, that EF is not a direct
measure of myocardial function, as it measures the cavity, not the
myocardial deformation. At best, it could be characterised as an
indirect measure. Does this matter? Yes. If we look at a few examples:
1: A person with an EDV of 125 ml, a stroke volume of 70 ml has an EF
of 56%, which is fairly normal values for a grown man.
2: A dilated ventricle to 250 ml, with maintained stroke volume of 70
ml, gives an EF of 28%, which is reduced. This is in accordance with
reduced systolic function.
3: Concentric hypertrophy reduced the cavity volume. A little old lady
of 80 years with concentric hypertrophy may have a cavity of 75 ml, a
stroke volume of 40 ml and an EF of 53%. Thus EF is normal, but in
terms of stroke volume, the systolic function is not!
Fractional shortening
As M-mode was the first echo modality, the Fractional shortening of the
LV cavity was the first LV systolic functional measure by echo. The
fractional shortening is defined as
FS
= (LVIDD - LVIDS)/LVIDD thus, in fact being an one-dimensional
version of EF. Diameter is conventionally measured to the endocardium,
so the fractional shortening is more precisely the endocardional
fractionla shortening. It's less accurate than the EF when there is
regional dysfunction, as the measured fractional shortening will be
generalised to the whole ventricle.
The relation between wall thickening and fractional shortening is
ilustrated below:
Thus, fractional shortening and wall thickening may be considered
inversely reklated. But they are not interchangeable as measures of
"radial function" as the same erroneous results will be obtained by the
fractional
shortening as of EF, as shown by the example
below.
Thus, the EF or FS is a measure that actually only works with dilation
of the
ventricles, and becomes erroneous in the cases of reduced EDV. Because
this has been poorly recognised, it has lead to some fairly
bizarre results. As systolic function has been measured by EF, and
diastolic function with mitral flow parameters, the hypothesis of
"isolated diastolic heart failure" has been proposed. At the outset,
measuring systolic and diastolic function by different measures with
different sensitivity, is methodological nonsense in any case.
This has been realised, ad the term is now substituted with the term
"Heart failure with normal ejection fraction" (
HFNEF).
But
as EF as a measure of systolic
function in the case of small,
hypertrophic ventricles is meaningless, the whole concept is
still
dubious.
The EF is introduced to characterise the reduced myocardial function in
dilation, by normalising an unchanged stroke volume for the
increasingly dilated ventricle. But this does not work the other way,
if the EDV does not increase or even decreases, the ratio has no
logical physiological meaning.
The erroneous comparison between longitudinal
strain and fractional shortening:
The
incompressibility
principle tells
us that as the wall shortens in the
longitudinal and circumferential direction, it has to thicken in the
transverse direction, and the relation is
geometrically
determined. Thus the longitudinal and transverse function as
measured by strain should be interrelated. Reports about radial
compensation of reduced longitudinal function is in direct opposition
to the incompressibility principle. The problem arises if we do
not measure the same values for longitudinal and radial function. It is
quite common to measure longitudinal strain, i.e. wall or segment
shortening as a measure of longitudinal function. On the other hand the
fractional shortening of the chamber diameter is a well established
measure of global and radial function. But in the case of hypertrophy,
this may lead to completely erroneous conclusions about the changes in
radial versus global function, as shown in the theoretical treatment
below.
In this theoretical M-mode of the LV, a
normal
ventricle has a wall thickness of 1 cm, an internal end diastolic
chamber diameter (EDD) of 4 cm, resulting in an external diameter of 6
cm. As most of the wall thickening is inward, with little change in
outward diameter (except in the case of differing filling pressures on
the two sides), an end systolic wall thickness of 1.5 cm will result in
a diameter shortening of 1 cm and an end systolic chamber diameter of 3
cm. Thus, wall thickening (WT, transmural strain) is (1.5 cm - 1 cm) /
1 cm = 50%, chamber diameter reduction is 1 cm, fractional shortening
(FS) is (4 cm - 3 cm) / 4 cm = 25%. In the case of concentric
hypertrophy, the chamber diameter is reduced due to increased wall
thickness. A hypertrophy leading to a wall thickness of 1.5 cm,
will give an EDD of 3 cm. A systolic wall thickening of 0.5 cm
will then be (2 cm - 1.5 cm) / 1.5 cm = 33%; i.e. a clear reduction in
radial function. But 1 cm diameter shortening is FS = (3 cm - 2
cm) / 3 cm = 33%, an apparent
increase in radial function, due to geometrical misconception!
From the reasoning above, any conclusions about radial function based
on fractional shortening in the presence of hypertrophy may be
erroneous, and the term radial function needs to be defined. The
conclusion that there is radial compensation for reduced longitudinal
function should be reserved to the cases where WT is increased.
It is extremely
important that if
longitudinal and "radial function" are compared, care should be taken
that the measurements are comparable. To compare for instance
fractional shortening of the LV diameter with longitudinal strain (wall
shortening), is comparing two different measures, and may lead to
completely erroneous conclusions as shown above, where fractional
shortening increases but wall thickening decreases.
Wall measurements - long axis systolic function.
Wall thickening is a measure of
systolic deformation. It can be
assessed semi quantitatively in B-mode.
Wall motion score index (WMSI) by B.mode, being the average
of
wall motion
score of all evaluable segments becomes a measure of
global
function, and has been shown to correlate with EF in infarcted
ventricles (
40).
It has also been shown to be similar in sensitivity to reduced function
(and infarct size) to global strain (
189).
However, the index is useless unless there is regional differences. Any
dilated cardiomyopathy will show hypokinesia in all segments, giving a
WMSI of 2, regardless of EF.
Wall thickening measured in M-mode, however, is only available in
limited segments, and can only be generalised to global measures if the
ventricle is symmetric. In addition, as discussed
above, the wall thickening
is mainly a function of the long axis shortening, due to the
incompressibility of the heart muscle.
The
systolic long axis function
is measured by any means of any longitudinal motion or deformation.
I.e. Long axis shortening measured by
mitral annulus
motion or
global strain, or shortening
velocity / rate by
mitral annulus velocity
or
global strain rate.
Mitral annular systolic displacement
Mitral annular systolic displacement or excursion (MAE), and mitral
annular systolic velocities, are measurements of total ventricular
shortening and shortening velocity:

|
|
Longitudinal
M-mode through the mitral
ring, displaying the displacement of the mitral ring. The total
systolic displacement (MAE; mitral annulus excursion) can be
measured. If the MAE is divided by the end diastolic length
of the ventricle (which, in fact is a spatial derivation), it will give
a measure of the strain of the wall. The global strain of the left
ventricle is an average of more points of the wall.The longitudinal strain during systole is
thus MAE /LD.
|
Pulsed tissue Doppler of the mitral
ring. These are the velocity traces of the longitudinal motion,
while dividing by the end diastolic length results in the Lagrangian
strain rate (Which is different from the Eulerian strain rate that is
customarily used in ultrasound. This is discussed below.
|
The annular measurements reflect the total shortening of the ventricle,
and are thus measures of
global
longitudinal function.
The annular The term mitral annular descent or
mitral annular excursion (MAE) (
31,
35,
37,
40)
should be used. Atrioventricular plane descent (AVPD) (
30,
32,
34,
36)
is
incorrect, as the term also comprises the tricuspid part, and while
tricuspid displacement and velocity can be measured (and is higher than
in the left ventricle) , it is usually measured only in one point, and
the relative weights for the measurements is unclear.
The longitudinal shortening has been shown to be very closely related
to ejection fraction when comparing different patients with normal or
reduced left ventricular function (
30,
31,
32,
34,
35,
36,
40,
64),
as illustrated below:
When the left
ventricle dilates, the volume increases, and the stroke volume can be
maintained by a smaller fraction (Ejection fraction) of the total (end
diastolic) volume. At the same time, the cross sectional area
increases, so the volume can be maintained by a smaller stroke
length.
The relation between MAE and EF has
shown a correlation of 80 - 90%. However, the relation only holds in
dilated ventricles. In normal ventricles, the MAE is related to the
stroke volume (
13,
59,
60,
116).
In left ventricular hypertrophy, the MAE is reduced despite preserved
EF, and there is no correlation (
190).
In addition, the MAE is reduced in ventricles with normal ejection
fraction , the so-called HFNEF (
191),
i.e. despite normal ejection fraction.
The annular displacement has been shown to be more sensitive than EF
in predicting events in heart failure (
36,
192)
and hypertension (
193), indicating
that it is a more precise measure of systolic function, that the cavity
measurements. This may be due to the shortcoming of EF in small
ventricles / hypertrophy. There is also a trend towards a better
correlation with infarct size than EF (
150).
Also, the MAE correlates better with BNP in heart failure, that the
fractional shortening (
204).
Thus, the MAE is a more all round
useful measure of longitudinal function than EF.
There has been some arguments for measuring MAE only during ejection,
i.e. excluding the isovolumic phases (
194).
The value will be a little lower, and the main advantage seems to be
that
post systolic shortening, not being part of
the systolic work, will be eliminated.
Systolic annular displacement.
There is a small shortening in the isovolumic contraction phase (IVC),
and post systolic motion (PSM) after AVC, so the
systolic MAE is lower than the total MAE.
However, the total shortening is probably related to the total
ventricular size. This means that small ventricles has a lower MAE,
even if similar in relation to the total length. This also means a
lower stroke volume, of course, from a smaller ventricle. So the
relation MAE x cross sectional area = SV still holds. However, this
means that some of the variations in MAE are due to heart size, not
heart function, which mans that the relation with heart function has a
reduced explained variance. Theoretically, this means that the annular
displacement should be
normalised for heart size,
which also is the case when using
global strain
instead, being relative shortening. This is definitely necessary in
children (
159,
214),
where the varation in heart size is great, the advantage in adults,
where variation in heart size is less (and less that the diffenence
betweeen normal and pathological) is not documented.
Where should measurements be done?
As the displacement is
higher in
the lateral than the medial, it is
evident that the measurements are different if different sites are
chosen. All studies have used the average of four points: septal and
lateral in the four chamber view, and anterior and inferior in the
two-chamber view. Thus the average is fairly robust, representing a
global average. However, the main reason for using four points would
be to reduce variability (which is reduced by about 25% by using four
points instead of one (
40). In
addition, regional differences due to regional dysfunction may be
evened out,, however, we found that ring motion was reduced in all
points in localised infarcts (
40).
Peak annular systolic velocity
Peak annular velocity occurs early in systole, and may be less load
dependent, as
maximum afterload is reached later in systole. Peak velocity is related
to acceleration, which is a direct measure of force, and thus to
contractility. However, peak velocity is
not load independent, as increased load will result in a
delayed and blunted development of force and velocity, as opposed to
the pressure/volume relation. In addition, as most pressure work is
done before ejection, the
pressure work will
not be measured.
Peak systolic velocity (S') has been validated as a measure of
systolic function (
37,
38,
39,
40).
The correlation with EF is weaker than for MAE, which is not
unsurprising, EF and MAE being
end systolic
measures, and as such measures of the total systolic work, S' is
peak systolic, measuring peak systolic
performance.
One of the main advantages of tissue velocities is that systolic and
diastolic function are measured by the
same method. From the beginning, systolic function by EF was compared
to diastolic function by mitral flow, equivalent to comparing apples
with bananas. This lead to the concept of pure diastolic dysfunction,
which has later been shown to be erroneous (
202).
The correlation between systolic function S' and diastolic function e*
was found in an early study to be 0.6 over a wider range of ventricular
function (
201), and in the HUNT
study (
165)with
a large number (N=1266) and limited to healthy subjects, the
correlation was found to be 0.59.
The correlation reflects among other things, the physiological
mechanism that much of the diastolic
recoil is due to elastic stored energy from systolic contraction
(restoring forces), but also, and most important:
that systolic and diastolic function are
closely related.
Age dependent peak systolic, early
and late diastolic velocity in normals from the HUNT study (165).
The early diastolic velocities are higher
than the systolic, and the decline is thus steeper, but the relation is
evident.
In another study (
202) it was found
that the systolic function by S' was reduced in patients with heart
failure with normal ejection fraction. This led to a renaming of the
state that up to then was called "diastolic heart failure" to "heart
failure with normal ejection fraction". This, of course corrects the
implied, but mistaken assumption that there existed a pure diastolic
failure. However, it does not address the fundamental problem, which is
one of methodology, that EF should not been used in normal sized or
smaller ventricles.
The S' has been shown to be sensitive for reduced function in relatives
who are mutation positive, of patients with manifest hypertrophic
cardiomyopathy, despite having normal EF and no hypertrophy (
203). The
diastolic function by tissue Doppler was
similarly decreased. It also correlates better with BNP in heart
failure than the fractional shortening (
204).
Thus, the peak systolic annular velocity is useful in that it is a
better marker of systolic function, and that it offers a measure that
allows direct comparison of systolic and diastolic function.
Where
should measurements be done?
As the velocities are higher in the lateral than the medial, it is
evident that the measurements are different if different sites are
chosen. This can be seen from the
HUNT
study (
165).
The initial studies (
37,
38,
39) used the average of four sites as
a measure of global systolic function. In the HUNT study, however,
there were no difference between the peak systolic velocity (S') mean
of lateral and septal, and the mean of all four points. However,
Thorstensen et al (
154) did show
that reproducibility was about 35% better using four point average
(p<0.001), in
line with what was found earlier (
40),
even if the mean values were the same.
Global strain and strain rate
Global strain and strain rate, may be taken as global measures of
ventricular function. This can be achieved simply by measuring and
averaging the strain/strain rate in all segments of the
ventricle.
However, there is one caveat:
Commercial software may give segmental values for six segments in
each
imaging plane, resulting in a total of 18 segmental values.
However,
this results in equal weight given to all myocardial levels, despite
there being much less myocardium in the apical level. In order
to ensure
that the average value gives similar weight to all parts of the
myocardium, only four segments in the apical level should be included,
as recommended by ASE/EAE (
146).
If
not, the global measures may be misleading. (This is doubtful in the
global
strain measurement by
2D strain).
The global strain by this application also is somewhat
processing dependent, as
discussed
below.
Strain and strain rate, however should
not be normalised for body size. Both measures are deformation
per length, i.e. in fact normalised already for the size of the
ventricle. Further normalisation for body size (which in fact is a
correlate of healthy heart size), will then be erroneous. This is
analogous to the fact that EF, which is stroke volume normalised to end
diastolic volume, is never normalised again for BSA.
Global strain by speckle tracking has been introduced as a new measure
of global left ventricular function (
147).
This compensates for the shortcomings of ejection fraction, being both
more correct in the case of small or hypertrophic ventricles, and more
sensitive (
149).
In the 2D strain
application, it should be noted that the application relies heavily on
the AV plane motion, and then distributes the motion along the wall as
explained and shown
above. By
this
method,
regional artifacts as drop outs and reverberations will have less
impact, which is an
advantage
in measuring global function. (As it may
be a
disadvantage in
regional function, as the same smoothing may
reduce
the sensitivity to regional reduced function).
It is unclear whether this application actually corrects for the
reduced
amount of myocardium in the apex, giving at the outset 6 segments per
view, or 18 segments in total. Bull's eye plots seem to show 17
segments, but whether this is carried over to the calculation of global
strain is uncertain.
Global longitudinal
strain by this method, has shown a trend to be more
sensitive to infarct
size and correlate better with infarct mass than EF.
Global longitudinal strain is
thus a measure of wall shortening, normalised for the length of the
wall, as length is measured along the curvature. Whether this allows
sufficiently for the reduced amount of myocardium in the apex, seems
unclear, as the referred study included 33 anterior and only 7 inferior
infarcts. Annulus displacement had a slightly less diagnostic accuracy
than global
strain, but whether this was significant is less clear.
Normalising the
annular displacement for LV length (see below), did not show
ovious improvement in diagnostic ability, in this group (
150).
However, annular displacement normalised for LV length
IS a measure of
longitudinal strain. Recent studies in children has shown normalised
displacement to be an age independent measure of systolic performance (
159,
214),
i.e. in the instance where the
variation in LV size is greatest in the normals.
Thus, it is emerging evidence that global strain, adds incremental
value to the simple AV-plane motion (
159).
This is credible, some of the variability in MAE will be due to
differences in LV size, and
normalising
will remove this variability and give a tighter relation to pumping
parameters normalised for body size, and thus a higher diagnostic
discriminatory value. This probably has most importance when normal
variations in body and heart size is biggest (as in children) and least
importance where normal variability is lower, and variation between
normal and pathological is great (as in dilated heart failure). None of
the methods for normalisation, however, have established superiority.
Normalised displacement and velocity.
As described
above,
the annular
displacement
divided by the length of the ventricle is a measure of global strain.
Similarly, the annular velocity divided by the ventricular length is a
measure of global strain rate. In fact, both displacement and velocity
can be normalised for (divided by) the distance from the apex to the
point of measurement as shown below.
Normalised
velocity/ displacement. The velocity and
displacement in each point along the wall from apex to base is the
resultant (sum) of the contraction of all the segments apical to the
measuring point. Thus dividing the velocity or displacement at a
certain point along the wall by the distance of the point from the
apex, will normalise the velocity and displacement for the distance,
resulting in values that are similar to strain rate and strain.
Mainly, this approach will be a compensation for the velocity / motion
differences between apex and base, making the evaluation of these
measures position independent (
54).
For displacement, being maximal at end systole, this will be similar to
the
strain of
the
whole myocardial segment between the apex and the measurement point.
The normalised annular displacement will be a measure of the global
strain, making it less dependent on ventricular size (and thus, body
size).
Recent studies in children has shown normalised
displacement to be an age independent measure of systolic performance (
159,
214),
i.e. in the instance where the
variation in LV size is greatest in the normals. The study in children (
159) did show better correlation with
EF over a wide range of pathology and age. In a small study in normal
adults, it has shown better correlation with EF (
217), which may be an indication that
it removes variability due to LV size. However, introducing another
measure (LV length) will increase the measurement variability of the
composite parameter, and thus, the advantege is still uncklear.
Thus, it is emerging evidence that normalisation of MAE, adds
incremental value to the simple AV-plane motion. This is credible, some
of the variability in MAE will be due to differences in LV size, and
normalising
will remove this variability and give a tighter relation to pumping
parameters normalised for body size, and thus a higher diagnostic
discriminatory value. This probably has most importance when normal
variations in body and heart size is biggest (as in children) and least
importance where normal variability is lower, and variation between
normal and pathological is great (as in dilated heart failure). None of
the methods for normalisation, however, have established superiority.
The normalised velocity may also be taken as a global strain rate.
This however, will be the Lagrangian strain
rate, not the Eulerian, with the differences described
above.
Peak annular velocity
normalised
for end diastolic LV length, will yield peak Lagrangian
strain rate. Peak annular velocity divided by the
length at the time of peak velocity, but this will be Eulerian
strain at the time of peak velocity, not peak Eulerian strain rate, and
thus be less meaningful. There has been no documentation of this so
far.
Normalisation of velocities is thus less established, and systolic
velocities
are related to diastolic velocities.
Peak systolic versus end systolic
measures of ventricular function.
Peak systolic measures are the measures of peak ventricular
performance, and can be measured as peak ejection velocity in the LVOT,
peak annular systolic velocity, and global ventricular strain rate.
These occur early in systole, and may be less load dependent, as
maximum afterload is reached later in systole. Peak velocity is related
to acceleration, which is a direct measure of force, and thus to
contractility. However, they are
not completely load independent, as increased load will result in a
delayed and blunted development of force and velocity, as opposed to
the pressure/volume relation.
End systolic measures on the other side, are measures of the total work
performed by the left ventricle during ejection. This is influenced not
only by force, but also by load (resistance), and the ejection time
(HR).
They are stroke volume, annular displacement and global strain, in
addition to EF. Whether this influences the sensitivity of the
measures, is not clear so far.
There is, however, little evidence directly comparing displacement /
strain to velocity / strain rate at varying load, and the few and small
studies that are published seems to indicate a very similar load
response.
Normal systolic values
Normal values are necessary if measurements are to be used
diagnostically. In addition, they will give additional information
about physiology. In the north Tröndelag population
(HUNT) study, 1266 subjects without known heart disease, hypertension
and diabetes were randomly selected from the total study population of
49 827, and subjects with clinically significant findings on
echocardiography (a total of only 30) were excluded. (
153)
This is the largest normal population study of echocardiographic strain
and strain rate rate to date. End systolic strain and peak
systolic strain rate was measured by the
combined
tissue
Doppler / speckle tracking segmental strain application of the
Norwegian University of Science and Technology, but the results were
compared to
other methods
in a subset of subjects, showing small differences.
The study consisted of
673 women with a mean BP of 127/71 ,mean age of 47,3 years and BMI of
25.8 and 623 men, with
mean BP of 133/77, mean age of 50.6 and BMI of 26.5. Both sexes were
normally distributed with an SD of 13.6 and 13.7 years, respectively.
20%
of both sexes were current smokers. Basic echo findings
are in accordance with other studies, like the findings
of Schirmer et al (
156,
157),
so the study population may
be assumed to be representative.
Normal values for systolic
velocities of the right and left ventricle from the
HUNT study.
|
Left
ventricle, mean
of 4 walls
|
Right
ventricle (free
wall)
|
|
S' (pw TDI)
|
S' cTDI
|
S' (pwTDI)
|
Females
|
|
|
|
< 40 years
|
8.9 (1.1)
|
7.2 ( 1.0)
|
13.0 (1.8)
|
40 - 60 years
|
8.1 (1.2)
|
6.5 (1.0)
|
12.4 (1.9)
|
> 60 years
|
7.2 (1.2)
|
5.7 (1.1)
|
11.8 (2.0) |
All
|
8.2 (1.3)
|
6.6 (1.1)
|
12.5 (1.9)
|
Males
|
|
|
|
< 40 years
|
9.4 (1.4)
|
7.6 (1.2)
|
13.2 (2.0)
|
40 - 60 years
|
8.6 (1.3)
|
6.9 (1.3)
|
12.8 (2.2)
|
> 60 years
|
8.0 (1.3)
|
6.4 (1.2)
|
12.5 (2.3)
|
All
|
8.6 (1.4)
|
6.9 (1.3)
|
12.8 (2.2)
|
Annular velocities by sex and
age. Values are mean (SD). pwTDI: Pulsed Tissue Doppler recorded at
the top of the spectrum with minimum gain, c TDI: colour TDI.
Normal
range is customary defined as mean ± 2 SD.
The study is based on 1266 healthy individuals from the HUNT
study by Dalen et al (
165).
The age
dependency of values is evident. Colour tissue Doppler gives mean
values, which are consistently lower than pulsed wave values, as
discussed
here.
It is evident that the systolic values
decline
with age, as does the early
diastolic.
Normal values for left ventricular
strain and strain
rate from the HUNT study
|
Female
|
Male
|
|
End
systolic
strain (%)
|
Peak
systolic strain rate
|
End
systolic
strain |
Peak
systolic strain rate |
< 40 years
|
-17.9% (2.1)
|
-1.09s-1
(0.12)
|
-16.8% (2.0)
|
-1.06s-1
(0.13) |
40 - 60 years
|
-17.6% (2.1)
|
-1.06s-1
(0.13) |
-18.8% (2.2)
|
-1.01s-1
(0.12) |
> 60 years
|
-15.9% (2.4)
|
-0.97s-1
(0.14) |
-15.5% (2.4)
|
-0.97s-1
(0.14) |
Over all
|
-17.4% (2.3)
|
-1.05s-1
(0.13) |
-15.9% (2.3)
|
-1.01s-1
(0.13) |
Values are given as mean ( SD). The
customary definition of normal
values as mean ± 2SD, giving
about 95% of the normal population, results in wider normal limits than
previously shown as cut off values in small patient studies. The values
were normally distributed, and with no clinically significant
differences between levels or walls. Values decline with age, as does
the velocity.
Regional systolic function
The regional systolic function is traditionally shown as wall motion
score:
- Normal
- hypokinetic
- Akinetic
- Dyskinetic
Wall motion score index (WMSI), being the average of wall motion
score of all evaluable segments becomes a measure of
global function, and has been
shown to correlate with EF in infarcted
ventricles (
40).
However, the index
is useless unless there is regional differences. Any dilated
cardiomyopathy will show hypokinesia in all segments, giving a WMSI of
2, regardless of EF. Thus, wall motion score is useful only in regional
dysfunction.
Segmental division of the left
ventricle. The segments are related to different vascular territories,
as shown by the colours. After Lang et al (146).
However, in the figure given in that paper, the apicolateral segment is
given as Cx or LAD, while the apical inferolateral is not, despite the
model is only giving four segements in the apex. Thus, there is a
slight inconsistency.
This segmental model gives a longitudinal resolution of the model of
about 3 cm, and a circumferential of 60°, which may be considered
low. However, in relation to vascular territories, it seems sufficient,
and deformation rate measurements with higher resolutions (which are
possible with both speckle tracking and tissue Doppler) have not so far
demonstrated added clinical value.
It is regional systolic dysfunction the deformation imaging has it's
main use, as it makes it possible to differentiate between passive
motion due to
tethering and active
contraction. Longitudinal strain can give the
wall
motion score by parametric imaging . It has been shown to give
about the same infrmatin as wall thickening by B-mode (
6,
7).
Segment interaction
The segment-segment interaction is
necessary to understand the effects of regional function measured by
deformation imaging.
Diagram of longitudinal segment
interaction. the longitudinal shortening of one segment results in
shortening of the segment itself (orange arrows), but also in motion
(red arrows) of the segments basally to it. (In this illustration, the
red arrows show the motion of the middle of the segment, meaning that
it also included the effect of the shortening of the apical half of the
segment itself.)and the motion of each
segment is equal to the summation of the shortening of the segments apically. However, the primary effect
is force generation. And this means that contraction in one segment
results in a force applied to the neighboring segments. This force has
different effects, as the apex is considered anchored (by the recoil
force), while the midwall segment has force applied from both
sides,
and the basal segment is freely movable. The main point is that
the force from neighboring segments may be considered part of the load
of each segment, and that motion iis secondary to deformation, but
deformation is secondary to force and load.
The
load dependency of deformation
parameters, as well as the understanding of load as partly the global
load (determined by the radius of curvature and the intracavitary
pressure), and the regional load, being dependent on the froce from
neighbouring segements, is the basis for both the changes in systolic
deformation and the
post systolic shortening. Thus
the main point is that deformation parameters
are load
dependent. But this means that if the contractility in one
segment is reduced, the part of the load of other segments that is
caused by the contraction from that segment, is reduced. But that
means that deformation by neighbouring segments may increase, due
to reduced load, and, concomitantly, the affected segment will show
reduced deformation. This again, is due to the point that the global
deformation happens within a framework of a virtual "
eggshell",
and
the AV plane. The global loss of
contractility by a regional process (as ischemia or infarction) will
reduce the global deformation, and within the ventricle the
regional
deformation will reflect the inequalities of force development. Thus,
regional loss of contractility may be
inferred
from the reduced regional deformation.
Both acute ischemia (
46,
99), as well as loss of longitudinal
fibres in myocardial infarction (
210)
will lead to loss of contractile force in the longitudinal direction.
This is equivalent to a local increase in the load/contractility
relation. This, however, may give different deformation patterns,
depending of the
amount and location of the contractility loss. Again, it is to be
emphasized that the deformation does
not measure
contractility directly, nor is deformation dependent on muscle
action alone.
 |

|
| Deformation patterns in apical
loss of contractility. A: normal pattern as in the diagram above. B: partial loss of
contractility, as shown by the shorter black arrow pulling on the
midwall segment. In this case several things may happen: If the
residual contractility is just about to balance the force from the two
other segments, no deformation occurs, thus the segment will be
akinetic, but not due to a total loss of force. Thus, akinesia does not
necessarily mean total loss of function. If the contractility is a
little better, there will be shortening, i. e hypokinesia. If the
contractility is a little too small, there will be stretching, i.e.
dyskinesia as in C. In the case of akinesia shown here, there will be a
little motion of the middle of the midwall segment, due to the
shortening of the apical part, but not much, and the basal segment will
have substantially reduced motion as well, despite both segments having
normal shortening. C: Total loss of contractility. (Of course, in
this case normal function of the midwall is improbable, this is just an
illustration of the mechanics. In this case, as the apex is
anchored, there will be stretching of the apical segment. The
midwall segment may then have no motion, as the stretching of the apex
and the shortening of the basal segment may cancel out, as depicted
here. Or there may be net motion in the apical direction, as the
stretching may require more force than moving the (more freely moving)
basal segment. Also, especially in infarction, the picture may be
complicated by fibrosis. Heavy fibrosis in scarring may render the
segment totally un-stretchable, thus mimicking situation B. |
Apical infarct. In LAD infarcts, the whole
of the apex is usually affected, the infarct sits as a "cap" over the
apex, although the extension towards the base may vary in the various
walls. Thus, the reasoning in the illustration to the left holds for
the whole ventricle.
|
In apical infarcts, some of the mechanics is thus determined by the
fact that the apex is anchored (by the recoil force) and does not move.
In basal inferior or inferolateral infarcts, the infarcted segments are
situated close to the more freely moving ring.
Thus, even with loss of contractility, there will be less load, so the
base can shorten, and even with total loss of contractility, the
segment will move. so the tendency to stretching is less, and even in
functionless infarcts there may be no or nearly no stretching. However,
this is not the only point. In LAD infarcts, the infarcted
segments are situated in the apex, as shown above left, meaning that
all
walls are affected, although the extension towards the base may vary,
so the wall are not affected to the same degree. In Infarcts of the RCA
or distal Cx, as well as isolated obtuse or diagonal infarcts, the
infarct does not extend around the whole circumference, and the effect
is more regional as shown below.

|

|

|
Inferior infarct. A: Normal
function. (The arrows indicating normal shortening are smaller, to give
room for the hyperkinesia in the infarct situation in B.) B: Total
loss of force in the basal segment. Even with total
loss of force, the segment can be pulled along, due to the tethering effect, and the fact that the mitral
ring, as opposed to the apex, is movable. A perfect example of this can
be seen here. Thus the probability of any
great degree of stretching is less probable. A small amount of
stretching my be present, depending on the interaction with the other
segments pulling on the mitral ring. There is thus no force from
the basal segment acting on the rest of the wall, and thus the load on
the two other segments are reduced, leading to increased
shortening, which may be interpreted as "compensatory
hyperkinesia". However, this follows as a function of reduced load, not
hyper function. AS the mitral ring moves around the whole of the
circumference, the shortening normally distributed to three segments,
in the infarcted wall is only distributed to two.
|
Inferior infarct. There is slight
stretching, but the main point is the fact that the infarct only
affects the base and midwall of the inferior wall, and the base of the
septum. Thus only the basal part of 1/3 of the circumference is
affected.
|
The AV plane. Not only is the segments
around the mitral ring closely bound together, thus excluding the
possibility of each segment moving independently, but the mitral ring
itself is part of the whole AV plane, consisting of the connected rings
of the pulmonary arteru (PA), the aorta (Ao), the MItral (MV) and
the tricuspid valves. Thus even the possibility of tilting of themitral
ring as each wall functions differently, is severely restricted.
|
Thus, it may seem that in apical
infarcts, there is more resustance to the normal segments, as the
infarcted segments are stretched, and thus, there is a slightly higher
load, while the basal infarcts, sitting at a moving ring, will offer
less resistance to the normal segments, allowing them to shorten more.

|

|
| Inferior infarct at day 1, showing akinesia
in the basal segment (yellow curve) and hyperkinesia in the apex (blue
curve). The hyperkiesia can be explained by the load reduction due to
the lack of force from the infarcted segment. (Image courtesy of Charlotte Björk
Ingul). |
The same patient at day 7. Function in the
basal segment (yellow curve) can be seen to be nearly normalised, and
the shortening of the apical segment (blue curve) is correspondingly
reduced. (Image courtesy of Charlotte Björk
Ingul).
|
Even so, the basal infarcts, sitting on a moving ring, will counteract
the tendency to stretching as well, the other basals egments will move
the ring.
But to complicate things, it must be emphasized that
the mitral ring is stiff, each segment
does not move independently.
It's easy, looking at one wall at a time to see only one mitral segment
at a time. It has been maintained by some (
32), that the
motion of the mitral ring is semi- regional, identifying the wall,
although not the segment affected, but this is not the case. It might
be conceivable that tilting of the mitral ring in response to different
actions of the different walls in regional dysfunction might lead to
tilting of the ring. But as shown above there is no isolated mitral
ring, the ring is simply part of the much bigger fibrous AV plane, and
thus the possibility of the ring tilting is restricted.
Thus it can not be inferred that only the
segment close to the infarct
can identify the affected wall. In a study of 19 infarct patients
versus
19 control subjects, in 2003 (
40), we
did show that. There mitral ring motion were reduced in infarct
patients compared to controls, and more reduced in anterior than in
inferior infarcts due to the difference in infarct size. Thus,
segmental reduced function will not cause the ring to lag in part of
the circumference, so much as the total ring motion will be reduced as
a function of the reduced total shortening force. This may explain why
the global strain is just as useful as regional strain in assessing the
infarct size (
205).
|
EF
(%)
|
WMSI
|
MAE
(mm)
|
S'
(cm/s pwTDI)
|
S' (cm/s cTDI)
|
Controls:
|
55
|
1
|
16
|
9.9
|
7.6
|
Inferior
infarcts:
|
45
|
1.4
|
13
|
8.0
|
5.0
|
Anterior
infarcts:
|
38
|
1.7
|
11
|
7.5
|
4.7
|
All differences between controls vs patients were significant, and also
inferior vs anterior except for S' (by both TDI methods).
What was more important was that the
variability
of measurements around the ring was the same in controls and patients,
and there were no differences between segments close and remote to the
infarct in the patients. (Not only no significant differences, there
were no differences at all). Thus the concept of the segmental motion
of the ring being a semi regional measure is wrong.
The most interesting finding was that while the infarct segments did
show significantly lower strain rate compared to remote segments,
despite the higher variability of strain rate measures. However, when
the strain rate of the three segments in the affected wall was
averaged, the differences disappeared, confirming the finding from the
mitral ring measurements that
the total wall
shortening is not a regional
measure.
Thus, the global systolic motion of the ring is a measure of the
infarct size (
32), being reduced in
proportion to the total amount of longitudinal fibre loss (
210).
Motion is global function,
only deformation can be regional.
As shown above, motion parameters will
thus always reflect global function, only deformation parameters can
show regional
function. This can be seen both in systolic and
diastolic function. The myocardium
moves within the stiff framework of the annular plane and the
"eggshell", but within this, there are differences in deformation, both
in amount and timing, which will lead to segments deforming
differentially.
Thus, as
deformation is a result of tension, or rather tension versus
load, strain does not measure function directly. But the effect of the
force from neighbouring segments is part of load. Taking
regional
function into the concept of load, deformation imaging can
be
used to
infer
force, or
at least inequalities in force development, as shown
below.
This means that regional deformation is closer to contractility than
global
measures, which are dependent on
absolute load. And that is the main point in
regional diagnosis.
In relaxation, this means that while protodiastolic elongation is mid
ventricular, it will result in elongation also in the base, early
relaxation has different timing of deformation in different levels, but
still results in an over all motion of elongation. Segements may
contract differentially, but this is not reflected in regional
differences of motion. Finally, global strain is simply global ring
motion normalised for LV size.
Post systolic shortening
Post systolic shortening (PSS) means that the segment continues
shortening after the aortic valve closure, often after a short
relaxation giving one or two peaks a systolic and a post systolic, or a
single peak after AVC as shown in the figure below, left. A small
amount of post systolic shortening may be present in up to 1/3 of
normal segments (
47),
but not more than ca 3%. Pathological strain is concomitant with
reduced
systolic
strain, and higher post systolic strain (in magnitude), as well as
later peak PSS. Post systolic shortening and post systolic thickening
are to some degree equivalent, due to the incompressibility as
discussed
above.
It is evident that in a segment being stretched in systole, if there is
any elasticity at all, the segment will recoil in diastole, i.e. as a
function of the elastic force stored in the segment. (also, if the
segment had not returned to the original shape, the whole heart would
have been turned inside out in the time of a few minutes. Thus, stretch
recoil is a mechanism for post systolic shortening. However, PSS can be
seen in segment that have systolic shortening as well, as shown
below. In ischemia, post systolic shortening
develops
before
there is systolic stretching (
46,
100
), i.e. while there still is systolic shortening as shown in the stress
example
below.
Thus, PSS can be present where there is systolic shortening as well,
and here the mechanism has to be different from the recoil. It has been
proposed that storing of elastic forces due to the interaction with
normal segments during systole maybe a mechanism, but it is difficult
to see how this can be the case, as elasticity is a function of
stretch. In ischemia, as well as in loss of function, both the rate of
shortening as well as the total shortening (i.e. strain rate and
strain) is reduced (
208,
209)as shown
above. But in isotonic twitches in
isolated muscle, this reflects the force development. Thus the
increased relative load in ischemic / infarcted segments is important,
in that it delays the force development. This means that the affected
segment is still shortening when neighbouring segments relax, resulting
in reduced load on the affected segment, which then can contract while
the other segments relax. When healthy myocardium relaxes, the
delayed relaxation of the pathologic segments will cause them to
shorten, as a function of the reduced force in normal segments. Thus,
the post systolic shortening is a function of the interaction between
segments. In this phase, there is pressure decay as well, decreasing
global load, further reducing the load on the affected segments. Thus,
post systolic shortening is mainly the reduced, but prolonged
contraction of a segment due to ischemia and / or relative load
increase in the early diastolic interaction with normally relaxing
segments.
Diagram of post systolic
shortening in an apical segment. In systole, there is reduced
contractility (force) in the apical segment, causing reduced shortening
compared to the other segments. In early diastole, there is no force
from the normal segments, as they now are elongating during relaxation.
(Elongating being the result of elastic recoil from the systolic
compression as discussed above).
The prolonged contraction (force development) in the affected segment,
is allowed to continue shortening as it is not counteracted, causing
further shortening during early diastole.
That segment interaction is a prerequisite for PSS, can be seen in the
example
below, where there is total ischemia,
and hence, no normal segments and (almost) no PSS in the ischemic
segments.
In ischemia, the cytoplasmic calcium transient is also reduced, leading
to more prolongation of the contraction, thus there may be more PSS in
ischemia than in old infarcts where the mechanism is mainly relative
load. This seems to be the case as the presence of PSS decreases in the
three months following the infarcts (
92).
The post systolic shortening was about the same in border zone segments
and infarct segments, despite infarct segments having lower absolute
value of peak systolic strain rate. The PSS diappeared in the border
zone segments in a week.
As seen by the colour M-mode
below,
the
presence of post systolic shortening in a segment, leads to a delay in
the onset of
segmental lengthening compared to the normal segments, so the finding
is equivalent to
the delayed compression/expansion crossover described by some authors (
186).
Looking at
Strain rate,
it is evident that any systolic stretch with early diastolic recoil
will show up as post systolic shortening. However, with strain, it is
useful to separate between systolic shortening followed by post
systolic shortening. This is better shown in the curves with strain,
but also in the colour M-mode of strain rate, which in addition gives
the extent of PSS.
 |
 |
 |
Normal strain rate curves. Note that there
is a little shortening of the lateral wall (cyan curve) after AVC
(green vertical line). This is normal.
|
Two different instances of post systolic
shortening. Apicolaterally, there is stretching and then recoil after
AVC (cyan curve). There is little indication of
active contraction at
all (except possibly a little overshoot, but that may be elastic).
However, the stretchability and recoil indicates that the tissue has
not lost it's elasticity. Apicoseptally there is systolic shortening
and then further post systolic shortening (yellow curve), which thus
has to be active. It also shows the mechanism for PSS to be different
than recoil. In fact, these curve is very similar to the
curves in
the original work of Tennant and Wiggers from 1935 (46).
|
Reduced systolic shortening and presence of
post systolic
shortening in the apical segment (cyan curve), with normal systolic
shortening and no post systolic shortening in the basal segment (yellow
curve).
|
The presence of post systolic shortening in the earliest phase of acute
ischemia, was demonstrated already by Tennant and Wiggers in their
experimental work in 1935 (
46),
although in the paper they chose not to discuss the phenomenon,
concentrating on the initial stretching and decrease in amplitude of
shortening, and the full dyskinetic pattern showing up after a minute
or so. It is rumored that they considered this an artifact, but the
phenomenon is clearly visible in the published traces. Post systolic
thickening in ischemia was shown in a case by Jamal et al in 1999 (
185), and demonstrated during
angioplasty by Kukulski et al (
100).
Decreasing systolic and increasing post systolic shortening with
increasing ischemia is demonstrated below. (In fact, there is a
striking similarity with the traces by Tennant and Wiggers in their
original paper. It was shown to be present in both is chemic and
scarred
myocardium (
47),
in about 75 to 80% of segments.
Development of apical ischemia
during
stress echo.; showing normal contraction at baseline, increased during
low dose (10 µg/kg/min, may be a biphasic contraction at 20 µg/kg/min, not very evident in this
animation, but may be better visualised by stopping and scrolling the
loop in the clinical situation.

|

|

|
Colour SRI M-modes from septum of the same
examination, showing clearly at 20 µg/kg/min the development of a
prolonged shortening period in the apex, but still systolic
shortening as well. During peak stress, there is virtually no systolic
shortening, only post systolic.
|
Strain curves at 20 µg/kg/min (top) and peak stress (bottom), showing systolic hypokinesia
at low dose with PSS and akinesia in septum / dyskinesia laterally
with PSS. Thus, PPSS are seen also with hypokineqsia, and is not only a
recoil after stretch.
|
Post systolic shortening has been proposed to an additional diagnostic
criterion for ischemia in stress echo (
113),
but other studies has not shown additional diagnostic value of this (
128). Two examples of systolic
dyskinesia with post systolic shortening in myiocardial infarction are
shown below.

|
 |
| Bull's eye and three dimensional
reconstructions of a ventricle in systole (top), showing an area of
dyskinesia (blue) in the apex, and diastole (bottom), showing a larger
area of post
systolic shortening (yellow). |
Bulls eye from systole and early diastole
(top, left) , below 3D reconstruction
(bottom, left) in systole and M-modes from all six walls (right),
showing an inferior
infarct with slight
dyskinesia and more extensive akinesia in systole and post systolic
shortening
in the infarcted
wall. |
In
the systolic images, the areas of dyskinesia are especially evident,
but as in the stress example above, areas around may be
hypokinetic (not as evident in the parametric images), but in the
diastolic images the PSS is seen to be fairly extensive, proving
that this is not purely recoil.
According to the description
above,
if all segments are pathological, the PSS should be less obvious or
even absent due to
the lack of interaction with normal segments as demonstrated below.
 |
 |
| Severe ischemia in all walls in
a patient
with severe three vessel disease (among other things stenosis left
main, occluded LAD filled from RDP, even with occluded RCA filled from
collaterals) . Visually, the most striking finding is fall in EF
with increasing stress. |
Strain rate colour M-mode. No
significant PSS can be seen (Except possibly apicolaterally). Thus at
fiorst glance, the M-mode looks normal, at least concerning synchrony.
|
 |
 |
| Strain rate curves (left) and strain
(right) of the
ventricle at
peak stress. Again, no
significant PSS can be seen (Except possibly apicolaterally),
demonstrating clearly that there are little PSS when there are no
segments with normal contraction-relaxation cycles.
The AVC is evident from the phono traces. The strain curves (Left,
bottom, shows delayed and prolonged shortening, but more or less in
all segments. This is equivalent with the balanced ischemia of
scintigraphy. |
Presence of PSS may give asynchrony between walls, where almost all of
the wall may be out of phase, even if there are gradients of ischemia
as shown below.
 |
 |

|
| Stress echocardiography with development of
ischemia in the inferolateral wall. At peak stress, the whole of the
wall can be seen to move paradoxically, moving inwards (and towards the
apex) after end of septal contraction. Again, in a clinical
situation, the interpretation can be facilitated by stopping and
scrolling. |
The velocity (motion) confirms the visual
impression, the whole inferolateral wall moves downwards in systole,
and upwards after end systole (Yellow and green curves), while the
septum shows normal apically directed velocities giving a total
asynchrony between the two walls. This asynchrony is also evident by
the curved M-mode, starting a the inferior base, going through the apex
and ending at the septal base.This might be due to both apical and
basal ischemia. |
 |

|
The strain curves below, separates the
effects of the segments, showing systolic dyskinesia
(lengthening) with some net post systolic shortening in addition
to the recoil in the base (yellow curve), and systolic hypokinesia in
the apical segment (green curve) with post systolic shortening,
compared to a fairly normal strain curve in the septum. Thus,
deformation imaging showing most severe ischemic reaction in the basal
part, giving highest probability of a Cx ischemia, which was confirmed
angiographically.
|
In this case, the tissue velocities are sufficient to detect the
presence of ischemia, but the deformation imaging shows the location
and extent of the ischemia, while velocities shows
asynchrony of the whole inferolateral wall.
Asynchrony:
The presence of regional systolic dysfunction in
combination with post systolic shortening, may cause asynchronous
motion of a whole wall, as shown above. This means that the presence of
asynchrony in motion imaging is not specific. A further example, also
from stress echo; i.e. ischemia, is shown below.

|

|
Velocity images showing motion towards the
apex in red, away from apex in blue. Left, systolic 3D
reconstructed image, showing normal motion in the septum and inferior
wall, and paradoxical motion in the inferolateral, lateral and anterior
wall. Right, om top are bull's eye from systole, showing the same, as
well as early diastole showing inverse motion during the e-phase, i. e
motion of the whole wall towards the apex in diastole. Apparently, the
whole anterolateral half of the ventricle is ischemic .
|
Strain rate images from the same recording,
left systole, right early diastole, showing that the ischemia is due to
a smaller ischemic area in the
inferolateral, lateral and anterior apex, where there is streching
during systole (blue). This stretching, results in the midwall
and basal
segments moving away from the apex, despite contracting normally. In
early diastole there is recoil in the ischemic area (yellow), resulting
in
anterior diastolic motion in the whole of the wall. In this case,
the ischemia is obviously limited to a part of the apex, the rest of
the motion abnormalities being due to tethering. |
In this case, it is evident that asynchrony between walls, as seeen by
motion imaging is not real asynchrony, but an effect of
tethering to a
smaller asynchronous area. Thus, simply showing asynchrony by motion
imaging is insufficient in the diagnosis of condiuction abnomalitiy
induced asynchrony.

|

|
Stress echo. In this case, the image is
suspect of a delayed inward motion of the base of the wall at start
systole at medium and peak dose.
|
In this peak stress image, the tissue
Doppler confirms the presence of initial asynchrony: The whole of the
inferolateral wall seem to show dyskinesia (Yellow and cyan curve),
with early motion (after IVC) away from apex. It seems to be most
pronounced in the apical part.The septal base moves normally, toward
apex (red curve). By placing the sample volume in the aortic ostium,
the high velocities of the aortic closure is identified, giving the
timing of end systole. (This can be more reliably done from basal
recordings as shown above, but these were
early days. Actually the AVC timing here is a little off, it should be
closer to 0.27 or 0.28). This timing is then transferred to the
deformation images to the left.
|

|

|
| Strain rate (left) and strain (right)
showing that there is a slight reversal of shortening in the early
diastole, but only in the base (cyan) . The delay, however, is
shown to be entirely within systole. The clinical meaning of this is
uncertain, but it was not due to ischemia, as the coronary angiography
was completely ("super"-) normal. |
In this case, the deformation imaging helps in the timing confirms the
dely,but shows it to be less pronounced, abnd not indicative of
ischemia, and it also helps in showing the the extent, compared to
tissue velocity.
It must be emphasized that the presence of PSS is mainly a measure of
inhomogeneity of force development, due to differences in activation,
load or contractility, and not as specific marker of ischemia.In
pathological myocardium this has also been demonstrated
in hypertrophic cardiomyopathy, where the prolonged contraction
persists into diastole, even causing ejection from the hypertrophied
apex
(
87).
|

|
| Recording from a patient with
apical hypertrophic cardiomyopathy. During systole there is virtual
obliteration of the apical cavity. Ejection can be seen in blue,
and there is
a delayed, separate ejection from the apex due to delayed relaxation.
There is an ordinary mitral inflow (red), but no filling of the apex in
the early phase (E-wave), while the late phase (A-wave) can be seen to
fill the apex. Left, a combined image in HPRF and
colour M-mode. The PRF is adjusted to place two samples at the
mitral annulus and in the mid ventricle just at the outlet of the apex.
The mitral filling is shown by the green
arrows, and the late filling of the apex is marked by the blue
arrow. In addition, there is a dynamic mid ventricular gradient
shown by the red arrow, with aliasing in the ejection signal in colour
Doppler. The delayed ejection from the apex is marked by the yellow
arrow (the case
is described in (87). |
 |
| Strain rate from the same patient, showing
PSS in the apical part of the septum, and nearly the whole of the
lateral wall. (images were made in a very early software, but yellow is
still shortening, blue lengthening.) |
Also in hypertension is PSS
a frequent finding (
187).
The new concept of "mechanical dispersion",
meaning unequal duration of the shortening phase, seems to be due to
variable amount of PSS within the ventricle.
Regional circumferential
strain
The regional circumferential strain may be affected in different ways,
depending on the degree of transmurality, as sub endocardial infarcts
to little degree affects the midwall circumferential fibres, with
little or no effect on circumferential force. Transmural infarcts
affects those, leading to regional loss of circumferential force.
However, as circumferential strain is defined as midwall
circumferential shortening, even sub endocardial infarcts will result
in loss of circumferential strain due to loss of substance. As
subendocardial infarcts leads to loss of subendocardial fibres, there
will be a reduced wall thickness and also less wall thickening in the
infarcted area in absolute terms. However, thickening in relative terms
(percent of end diastolic wall thickness. i.e. strain) may be less
affected (i.e. less absolute thickening in a thinner wall may give the
same strain).
However, as the wall is thinner, the midwall circumference is
greater, but shortens less (in absolute terms) due to reduced wall
thickening. Thus, there will be
less circumferential
shortening of a
longer
circumference, i.e. the circumferential strain will be reduced, even
without affection of circumferential fibres as shown below. Howeever,
this will not lead to changed function by segment interaction, as the
circumferential force is not affected.
.
Diagram of a sub endocardial
infarct, not affecting circumferential fibres. The infarcted area is
depicted as an area with loss of
endocardial longitudinal fibres. As shown this leads to a thinner wall
in the infarcted area, and the midwall line is shifted outwards. As the
wall thickens in systole, the thickening is less in the infarcted area,
and this effect is even greater than the loss of fibres, as the effect
of the fibre rearrangement would have been most pronounced in the
endocardial layer due to the geometry. However, the effect on
transmural strain is unpredictable, as there is less thickening of a
thinner wall, the transmural strain, being relative, may be
preserved. The midwall circumference, however, is increased,
while the circumferential shortening (absolute) is decreased due to
less thickening as shown above. This
means less shortening of a longer
circumference. Thus, even without loss of
circumferential fibres, there has to be reduced circumferential strain
in the infarcted area. However, analysing program using an ROI, will
give the shortening of the midwall line of the ROI as circumferential
shortening. Thus, the ROI width, not the wall thickness is determining
factor, and this effect may be lost, if the ROI is not very well fitted
to the endocardium. This is indicated with the dotted lines
representing an assumed equal width ROI around the circumference.
The effect on the width of the ROI in 2D strain is illustrated below.
In infarcts affecting the circumferential fibres; i.e. infarcts with a
higher degree of transmurality, the geometry may be somewhat different.
In this case the circumferential force is reduced or absent in the
infarct, resulting an an imbalance of forces equivalent to what is seen
in the longitudinal direction. This means that there may be a systolic
stretching of the infarcted wall in the circumferential direction,
and also an increased circumferential shortening of the intact part of
the wall. This is shown below.
Exaggerated diagram of an
infarct affecting circumferential fibres. In this case, the force
(black arrows) from the circumferential fibres will pull and extend the
affected area in systole (boundary marked by green arrows), the infarct
being without balancing forces. The intraventricular pressure
will still serve to keep the wall distended (blue arrow), preventing
the infarcted area to be pulled straight across the gap. Thus, the wall
structures in the infarct may even diverge from each other, which is
hypothetically detectable in speckle tracking. Transmural strain may
also be more affected, as there will be no wall thickening due to
crowding and inward motion of longitudinal fibres (even if there are
some left).
Some authors have found a higher sensitivity for transmurality
by circumferential strain (221).
However, this may also be a function of reduced sensitivity for
subendocardial infarcts, due to the ROI issue as discussed above. The
effect of ROI width is illustrated below.