Faculty of Medicine and Health Sciences
Department of Circulation and medical imaging
Faculty of Medicine and Health Sciences
Department of Circulation and medical imaging
by
Asbjørn Støylen, Professor, Dr. Med
asbjorn.stoylen@ntnu.no
Deals with the basic physiological concepts; contractility and load, and the relation of deformation parameters to these, diastolic function, event timing and relation to valve function and intraventricular flow and pressures. Any imaging method deals with myocyte shortening, myocyte shortening is always a function of tension versus load, and thus any imaging method and measure is load dependent.
There are two methods for ultrasound deformation imaging, tissue Doppler and speckle tracking. Both methods have advantages and disadvantages, which are dealt with in the ultrasound chapters, but the basic physiology is method independent, so the basic physiology in this chapter is valid for both methods.
This section deals with the basic physiology as seen with various echo methods, and replaces most of
In the physiological aspects, as these were largely overlapping anyway.
There will be discussions of validity on composite measures, these will be based almost solely on physiological principles.
Of course the opposite caveat is also true. Biological plausibility is not evidence of effect, it's at bes hypothesis generating.
The first temporal derivatives of these parameters are
The spatial derivative of flow is
This is
Systolic deformation is the result of fibre shortening, although in a complex manner.
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Image of beating isolated myocyte, prepared so the cell fluoresces with the presence of free calcium in the cytoplasm, the cell is stimulated to generate action potential. In systole the cell can be seen to increase free calcium and simultanously shorten. In the cellular diastole, the cell can be seen to elongate, and simultaneously free calcium disappears from the cytoplasm. The isolated myocyte is the completely unloaded situation, where myocyte tension results directly in shortening, and where shortening is a direct measure of contraction.. Image courtesy of Ph.D. Tomas Stølen, cardiac exercise research group (CERG), Dept. of Circulation and Medical Imaging, Norwegian University of Science and technology. | Excitation-tension diagram. The Action potential triggers the influx of calcium, which triggers further release of Ca2+from sarcoplasmatic reticulum. Calcium binds to troponin, and allows activated (by ATP) myosin heads to bind to troponin sites on actin (cross bridge forming) and release energy, causing the filaments to slide along each other, as long as there is a high calcium concentration in the cytoplasm. As the cell membrane repolarised, this triggers the removal of calcium from the cytoplasm, mainly by the SERCA pumping it into the sarcoplasmatic reticulum again. The removal of free calcium is an energy (ATP) demanding, active transport of calcium into the sarcoplasmatic reticulum by SERCA.Thus, obviusly, both contraction and relaxation are ATP demanding processes, and energy depletion will affect both. |
Fibre shortening is not the same as contraction. We have both isometric and isotonic contraction:
Diagram of mucle sarcomere shortening. Shortening is the result of tension versus load. If the load is higher than the maximal tension the muscle can develop, the muscle will not shorten at all. Actin is still moved along the myosin , but the energy is stored as deformation within the sarcomere without shortening aof the sarcomere. This means that the force generated by contraction is stored as elastic tension in the muscle, and the contraction is isometric The middle figure, retaining the length of the baseline - left). If the load is less than the maximal tension, the muscle will start to shorten when tension equals load, and from there the contraction is isotonic - shortening at constant load. The right sarcomere is shorter than the baseline.
In physiology and pathophysiology, the main object in characterising increased or decreased muscle function is the contractility, which is the ability to develop tension independent of load. As imaging only measures length, the tension must be inferred from knowledge of load. Shortening is the result of tension vs. load as discussed later.
As the muscle uses some time to develop full tension, in most situations there will be a period before tension reaches load (an isometric phase) before tension equals load and the muscle starts to shorten (isotonically) (78).
The concept of load is important for understanding the contractility. Contractility is defined as the inherent capacity of the myocardium to contract independently of changes in the preload or afterload (79).
The preload is defined as the load present before contraction starts (79). This is illustrated below left.
Frank-Starlings law of the heart states: Dilatation is caused when increased venous return or decreased ejection increases end-diastolic volume. This form of dilatation, within physiological limits, increases the heart’s ability to do work. The stroke volume of the heart increases in response to an increase in the the end diastolic volume, when all other factors remain constant. (83)
Preload is the force acting on the muscle before contraction starts. Thus, it will stretch the muscle, and induce an increase in muscle tension. This is best explained in isometric experiments, where an isolated muscle is fixed in both ends , and tension is measured by a tensiometer.
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Isometric experiment. Isolated muscle with fixed length, tension measured by a tensiometer. The three measures of tension are the peak rate of force (tension) development, time to peak tension and peak tension. They are all measures of contractile function. | Adding a weight to the muscle without stimulating contraction, will stretch the muscle, before fixating the muscle and adding a tensiometer. This can be achieved without the weight, of course, simply by fixing the muscle at different lengths (pre stretch). |
With increasing length by pre stretch, the stretch will both increase passive tension (elasticity), but also increase the active tension developed when stimulated (78).
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Stretching the muscle before stimulation, increases tension. The increase in passive tension will be present at rest, before twitch. During twitch, there is an increase in total tension with increasing pre twitch length. The increase in contractile tension is then the difference between the passive and the total curve. | Isometric twitces with increasing pre-twitch length. In can be seen that as opposed to inotropy, time to peak tension do not increase, even though peak tension does, and, as a consequence of this the rate of force development (as the rise to higher peak during the same time gives higher rate). |
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Hypothetical length tension diagram, based on the sarcomere hypothesis, that by increased fibre length initially will increase the overlap between the myosin head regions and the troponin regions on actin, optimising the number of cross bridges that can be formed, and thus the peak tension obtainable. In this model there is an optimal length, then the available number of cross bridges, and thus the peak tension decline again. | Both shortening and shortening velocity can be seen to increase with increasing preload. |
Afterload is force added to the preload as resistance to the muscle shortening. Total load is preload + afterload. This is the force the muscle must overcome ( e. the tension the muscle must develop) in order to shorten, also termed wall stress. This can be expressed as wall stress, the force acting on the wall. This is proportional to both the intracavitary pressure, and radius. Wall stress, however is the tension per cross sectional thickness of the wall, i.e. if the same load is applied to different wall thicknesses, the thickest wall has lowest wall stress per mm thickness.
This is summed up in Laplaces law: Wall stress () is proportional to pressure (P) x radius (r), and inversely proportional to the wall thickness (h):
In an afterloaded contraction, the muscle must first build up force corresponding to the total load, before it can start shortening. When the force equals the load, further contraction is translated into shortening without tension increase (isotonic contraction). Thus, neither peak force nor time to peak force are relevant measures of contractility.
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The difference between pre- and afterload is illustrated here. After preload is added, a support is placed, preventing further stretch of the muscle when another weight is added. This second weight is the afterload. When the muscle contracts, it has to develop a tension that is equal to the total load, before it can shorten. If the peak force is higher than the total load, the muscle will then shorten without generating more tension, in an isotoniccontraction. | Isotonic isometric twitches tension diagrams above, length diagrams below. From the diagrams, it is evident that shortening only starts after tension have reached load, and then, the tension is constant while the muscle shortens. Thus the first part of an unloaded contraction is also shortening, while the first part of a loaded contraction is isometric, becoming isotonic after tension equals load. Peak rate of force generation (RFD), occurs during the isometric phase (except in the unloaded phase, where there is no tension development). Peak rate on the other hand occurs during the first part of the shortening, after tension = load, and is thus later than peak RFD. The figure also shows that shortening decreases as load increases, as more of the total work is taken up in tension development. This experiment shows that both peak rate of shortening and peak shortening, i.e. peak strain rate are affected by afterload. Both peak shortening and peak rate of shortening are affected by pre- and afterload.The curves are explained further below. |
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Length tension diagram of a muscle twitch in an isolated muscle preparation. The muscle takes some time to develop the tension that equals the load, and during that period the contraction is isometric, with no shortening. Shortening starts when tension equals load. When the muscle relaxes, relaxation induces shortening until tension again equals load, after that relaxation is isometric. | Series of twitches with different loads. All twitches follow the same tension curve, i.e. shows the same contractility, but as load increases, shortening starts at later time points, and the shortening time as well as the extent and rate of shortening decrease. | Series of twitches with the same load, but with different contractility (ability to develop tension). With decreasing contractility, it takes longer to develop tension = load, the period of shortening as well as the extent and rate of shortening decrease. |
Looking at the whole heart cycle, there is a complex interaction between tension increase and devolution, and shortening and elongation whioch is discussed here.
Contractility is defined as the inherent capacity of the myocardium to contract independently of changes in the preload or afterload (79). The contractility decreases in myocardial failure and with betablocker, but increases with inotropic stimulation, such as adrenaline, noradrenaline or calcium.
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Shortening velocity and total shortening, Relation to preload and total load. Both shortening and velocity can be seen to decrease with increasing afterload (total load), but increase with preload. | Shortening velocity and total shortening, Relation to total load and inotropy. Both can be seen to increase with inotropy, but decrease with load. |
As we see, both increased preload and inotropy wil increase the muscles ability to develop tension, and thus the shortening and shortening velocity in an isotonic experiment. This can also be shown alternatively:
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Stretching the muscle before stimulation, increases tension. The increase in passive tension will be present at rest, before twitch, and is equal in baseline and inotropic state. During twitch, there is an increase in total tension with increasing pre twitch length. The increase in contractile tension is then the difference between the passive and the total curve. At a certain length, active tension starts to decline, even if passive and total tension still increases. This effect is additional to the effect of inotropy. | Myocardial shortening vs pre-, afterload and contractility. shortening increases with preload, as shown in both panels, although to a certain extent, until the preload insensitive zone (where active tension starts to decline, but passive tension still increases). Shortening is the resultant of force vs. afterload, the higher the afterload, the less the shortening, for a given contractility. Contractility is the load independent part of force development. The higher the contractility, the more the shortening for a given afterload. |
Thus, both shortening (strain) and peak rate of shortening (peak strain rate) is load dependent, and not independent measures of contractility, and looking at imaging, it is difficult to discern between reduced contractility and reduced load as shown below. This means, of course, that both strain and strain rate are load dependent.
Cardiac volumes
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Isotonic isometric twitches tension diagrams above, length diagrams below. From the diagrams, it is evident that shortening only starts after tension equals load, and thenthe muscle shortens isotonically. Peak rate of shortening (peak strain rate) is at the start of shortening, and then declines (the slope of the shortening curve) . Only in the totally unloaded situation does the muscle start to shorten art start contraction. In the loaded situations, peak rate of force generation (RFD), occurs during the isometric phase, before peak rate of shortening. Peak shortening (peak strain), on the other hand, is at the end of the isotonic phase. | Comparing a tension length diagram of an isotonic/isometric twitch, and a pressure/volume (Wiggers)diagram. I've added the division of pre ejection into protosystole and IVC as discused above. The ejection period is not isotonic, as pressure increases and then decreases, and the myocardial tension must follow a similar course. Thus the tension increase is only during the first part of ejection, and then tension decline, so last part of ejection is relaxation. However, the volume curve will reflect the fibre lengthening and shortening, which makes is very obvious that strain and strain rate are about volume changes. Peak shortening is at en ejection, when volume is smallest. The conventional Wiggers diagram as shown by Brutsaert (black) describe the steepest volume decrease at the time of AVO, but as flow takes some time to accelerate, the peak volume decrease rate (which equals peak flow rate, must be slightly delayed after AVO). This is shown by the red part of volume curve. |
Volume changes and muscle fibre length changes are analogous, of course, as systolic decrease in volume has to be accompanied by muscle shortening, diastolic increase in volume by muscle lengthening. As described in the basic concepts page, there is systolic circumferential and longitudinal shortening, and fibre shortening, while transmural strain is not fibre shortening, but thickening.
Deformation of the myocardium. In systole, there is volume reduction, as well as longitudinal and circumferential shortening. The strains are related to volume change, as shown above.
However, the strains are coordinates of the myocardial deformation, not direct measures of fibre shortening. While circumferential shortening is geometry related, longitudinal shortening is the one that is most closely related to volume change, and thus to fibre shortening.
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Shortening curves related to afterload, modified from the figure above. The shortening in percent, is equivalent to the longitudinal strain of the muscle. | Strain curve from a normal subject. The strain curve is fairly similar to the shortening curves to the left. | The picture shows a detailed LV volume curve from a healthy person by MUGA scintigraphy, showing how analoguous the strain curve is to the volume curve. |
So now, we can relate the physiological findings in isolated muscle to the volume changes originally descrtibed in Frank Starlings law:
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Acute increases in end diastolic filling, will increase the stroke volume along the curve shown. This effect was observed with both increased venous pressure, but also with decreased stroke volume in previous beat, resulting in an increased EDV. Within physiological limits there is an increase, but with increasing dilatation, there will be less response. However, at least in normal hearts, there is little evidence for a descending limb of the curves. It is important that Frank Starlings law is about acute volume changes, not chronic changes as in hypertrophy or dilation. The curve shows the preload (which is | SV versus afterload. Shortening is the resultant of force vs. afterload, the higher the afterload, the less the SV, for all preloads. |
Thus, global strain rate and strain are not load independent, as explained above. One would almost say of course. Force is the primary effect of contraction. Deformation is secondary to force, and depends on load. Motion is the summation of deformation. The systolic volume change of the ventricle is related to the resistance, which again is a function of both pressure and vascular resistance. What we measure with deformation parameters, is only the changes in volume, and thus the wall deformations.
thus, physiologically the rate and amount of longitudinal shortening should decrease with increasing load. Anything else would be counter intuitive.
The heart cycle can basically be described in terms of volume changes, which in turn are the function of ejection and filling:
This is slightly simplified, as will be explained later.
The true isovolumic contraction time (IVC) is defined from MVC to the start of ejection. Thus, in this phase there is no volume change, and, hence, no deformation.
This phase it on the other hand, the period of most rapid pressure rise, peak dP/dt, which occurs during IVC, close to the AVC (105). This represents the most rapid rate of force development (RFD). As this phase is basically not influenced by aortic pressure, it is largely afterload independent, except for the Laplac e effect: It's important to realise that the Frank-Starling balance and the pressure volume loop both relate to acute changes in ventricular volume. The physiologic differences in EDV between individuals, and in chronic LV enlargement, are not responsible for preload. On the other hand, by the law of Laplace:
LVEDV is part of the afterload, the tension must be proportional to the radius: As the intraventricular pressure acts on a larger surface, and thus the force that has to be generated must be proportional to the surface area for the same pressure. dP/dt increases with inotropy (106, 107)
On the other hand, dP/dt is preload dependent as any contraction measure (106, 107).
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Muscle twitches, showing that the most rapid rise of tension occurs during isometric contraction, but the highest rate of shortening occurs in the bunloaded muscle before peak RFD. | Peak rate of pressure rise, which is the closes correlate to the rate of force/tension development. This occurs during IVC |
Compliance (distensibility) is defined as C = 
V / P, i.e. how much does the volume increase for a given increment in pressure. Elastance is the inverse value E = P / V, i.e. how much does the pressure increase for a given increment in volume.
Compliance describes the volume a a function of pressure, and hence, should ideally be described in a volume-pressure diagram. The figure shows linear compliance, as well as decreasing compliance (decreasing volume per pressure increment), as can be seen in an elastic system. | Elastance, despite simply being the inverse of compliance, describes pressure as a function of the volume, and thus is best describel in a pressure-volume diagram. Here is shown linear elastance, as well asincreasing pressure for the same volume increments, as can be seen in an elastic system. | The pressure volume volum loops is generally described in a pressure volume diagram, i.e. showing volume as ordinate and pressure as abscissa. However, both compliance and elastance can be shown in this diagram, obviously, as both describe the ratios between pressure and volume increments. |
The length - tension relation in isolated muscle is equivalent to the pressure - volume relation in the intact heart. To go into the concept of cardiac work and contractility, the whole heart cycle has to be considered. Basically, pressure-volume loops are useful concepts for visualising and explaining the relations between stroke volume, pressure, contractility and LV work.
The pressure-volume loop on the right is derived from the pressure and volume curves on the left. It is basically a tool to assess and explain the relations between contractility, load and output. Isovolumic contraction (IVC) is between MVC and AVO, there being no volume change, this is a true isometric contraction. Isovolumic relaxation is between AVC and MVO, this is the isometric phase of relaxation. Pressure volume diagram of a heart where pressure is plotted against volume, a pressure-Volume loop (PV-loop). Pressure-volume loops are useful concepts for visualising and explaining the relations between stroke volume, pressure, contractility and LV work.The version shows the cardiac phases, with the same simplifications as the Wigger's diagram to the left; rapid filling is shown as a passive event, and as valve openings and closures are shown at pressure crossover. Time runs around the loop in a counterclockwise direction. The width of the loop is equal to the stroke volume. The top of the diagram shows the pressure curve during ejection, and the height of the loop is the systolic - diastolic pressure difference. The tangent to the end systolic pressure-volume, is the ventricular elastance 
P / V.
In filling of any hollow organ, the elastance is usually taken to mean how much pressure increase a given volume increment generates as described above, and the definition of elastance is P / V.
End systolic LV elastance is defined as the slope of the straight line trough the end systolic pressure volume corner of the pressure volume loop. It has been shown to be nearly linear trough multiple pressure volume loops obtained by manipulating the pre and afterload. However, the straight line has not been shown to be crossing the zero point, i.e. the volume is not necessary zero when pressure is zero.
End systolic elastance is a function of LV emptying. Thus this must be taken to mean how much pressure the LV must generate for a given volume ejection (SV), but the definition is P / V, and thus conforms with the definition of elastance.
The PV loop shows the relation between load, SV and contractility in acute changes.
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Effect of preload. Increased preload (increased LVEDV - the right side of the curve moves right, the loop becomes wider), will, through the Frank-Starling balance increase stroke volume. This increased stroke volume will be ejected at the same pressure, thus returning to the same point on the ESPVR line. | Effect of afterload. Increased afterload (increased SBP moving the top of the curve upwards), will reduce the stroke volume. The end systolic point moves up the ESPVR line, shortening the width of the loop, i.e reduced SV. | Effect of inotropy. Inotropy shifts the ESPVR line to the left, thus increasing the force and LV emptying, increasing stroke volume through reduced LVESV, but also increasing the pressure, both through increased contractile force and increased volume being ejected into the vascular bed. |
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In reality, an increased SV will cause increase in SBP, causing an increased afterload on the same beat, thus reducing the effect on SV somewhat, through interaction between pre- and afterload. | In reality, an acute increase in afterload, will reduce emptying (increased LVEDV), so on the next beat, the preload is increased, partly offsetting the effect of afterload on SV. | In reality, decreased LVESV, without increased venous return, will in the next beat result in reduced LVEDV, thus offsetting the effect of inotropy somewhat by reduced preload. |
Ventricular elastance, is thus a load independent measure of contractile force (84).
In reality, this index is not easily obtainable in the clinic, even in continuous invasive monitoring, as volume measurements are not available in routine monitoring. But in animal experiments, using conductance catheters, where multiple pre- and afterload manipulations can be done, and where the ESPVR can be obtained by linear regression, it serves as a reference method, to test other contractility indices.
However, the LV elastance may not be a perfect gold standard anyway.
- Firstly, using end ejection for end systole, means that measurements are done at a point in time where the myocardium is in relaxation (but not relaxed) phase as discussed below.
Several studies have found that the ESPVR is not a straight line, curvilinear depending on contractility (85, 86), and afterload (87, 88).
Peak systolic pressure volume relation might be closer to the real thing, but is not easily discernible.
- Secondly, as we'll see later, the true end systolic volume is not easily defined due to the protodiastolic volume decrease as discussed below.
Finally, however, of course these volume considerations are only related to acute changes. In inter individual differences in healthy individuals, as well as in LV dilation, the EDV is not a measure of preload, and the PV-loops can only be interpreted in relation to the individual. PCWP, on the other hand, is a measure of preload, that is relatively standardised and thus a more universal measure of preload when applied across individuals, although it does not take the Laplace effect into consideration, the LV volume will stioll contribute to after- (or total load).
Volume: It's important to realise that the Frank-Starling balance and the pressure volume loop both relate to acute changes in ventricular volume. The physiologic differences in EDV between individuals, and in chronic LV enlargement, are not responsible for preload. On the other hand, by the law of Laplace:
LVEDV is part of the afterload, the tension must be proportional to the radius: As the intraventricular pressure acts on a larger surface, and thus the force that has to be generated must be proportional to the surface area for the same pressure.
Pressure: The pressure, is the systolic pressure, which varies during the ejection:
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The ejection period is not isotonic, as pressure increases and then decreases, and the myocardial tension must follow a similar course. Thus the tension increase is only during the first part of ejection, and then tension decline so last part of ejection is relaxation. The volume curve (as given by Brutsaert - black) is erroneous. Due to the acceleration of blood, the peak flow rate, and thus the rate of volume decrease must be somewhat later than the AVO, as shown by the red curve. | With conventional pressure/flow recordings, the peak pressure / tension is around mid ejection, but looking at flow, peak flow through aortic ostium is much earlier during ejection. As flow rate equals the rate of volume decrease, peak flow must indicate peak rate of volume decrease. |
2. The tension increase is closely related to distension of the large vessels. The elasticity (compliance) of the arterial (especially the aortic) wall. During ejection, the volume ejected into the aorta and distends it. The more distensible the aortic wall (the higher the compliance V / P - the volume increase per pressure unit), the less the pressure in the aorta will rise, and the lower the CAP. The stiffer the aorta, the less it will be distended (the less the compliance) for a given pressure increase, or conversely the more the pressure has to be increased in order to inject a certain volume (stroke volume) into the aorta. Thus, increased arterial stiffness will increase the systolic pressure, and hence, the afterload (108).
During ejection, the volume is ejected into the aorta, which is distended, storing energy in the elastic properties, and then contracting again during diastole, acting as a diastolic pump, with energy stored from systole.
The ventricle ejects the full stroke volume into the arterial bed during systole. However, there is systolic run off to the periphery, so the volume distending the arterial bed is less then the stroke volume. Thus the V distending the artery is less than the total SV, and the difference is determined by the peripheral resistance, and the pressure drops toward end systole.
Ventricular ejection distends the arterial bed, and the elastic properties results in pressure increase related to the volume, and the compliance decreases as the volume increases. However, the run-off into the arterial bed, during systole, leads to the end systolic volume (and pressure) is less than the total SV. This run off is determined by the peripheral resistance.
The pressure is inversely related to the arterial compliance, or directly related to the elastance P / V (the pressure increase per volume unit), and is the inverse of arterial compliance. As the pre arteriolar arteriaø bed is elastic, there is thus no reason to believe that the arterial elastance is linear. Arterial eleastance could also be both end diastolic, peak systolic or end systolic.
Thus, what is called "effectve arterial elastance" is defined as SV / ESP. But this, of course defines a straight line slope, even if this is improbable, so this concept is only valid for end systolic pressure and SV. However, using ESP takes into account the whole of the SV, but also the fall i pressure from peak to end systole, which is a function of the run off during systole, and thus peripheral resistance. Thus, it is a measure of afterload taking both components of the arterial bed into account (260 - 262).
PV-loop illustrating effective arterial elastance. The arterial elastance is a measure of how much pressure the stroke volume generates in the arterial bed,
and is simply EA = ESP/SV, which is the slope of the line crossing zero at the end diastolic volume, and the point defined by the end systolic volume and pressure. It is the arterial pressure increase in relation to the stroke volume. The ventricular elastance is LV end systolic elastance, the pressure generated by the ventricle fpr ejecting the stroke volume; EES = ESP/ESV. But then, given that the theoretical volume of the LV for pressure P =0 is V = 0, then EA / EES = (ESP/SV)/(ESP/ESV) = ESV/SV = (EDV - SV)/SV = EDV/SV - 1 = 1/EF - 1. Thi9s of course is not necessarily the case, but depends on the placement and slop of the LVESPR line. Thus the concept of VA coupling simply eliminates the pressure, and ends up with a load dependent measure of LV performance as a result of this load!
Pulse wave reflection at the level of the aortic root. In the upper panel, pulse wave propagation velocity is lower, and the reflected wave from the previous heartbeat arrives after AVC, and will then augment the diastolic pressure only, not contributing to afterload. In the lower panel, due to a higher PWP, the reflected wave arrives before the AVC, and will augment the systolic pressure, which willl be higher than the peripheral pressure. Systolic pressure augmentation will thus add to the afterload.
Finally, of course, the body regulates both the cardiac performance and load in relation to the needs of the body:
Thus, of course, all factors of cardiac performance in the intact body may change in relation to the body's needs. But this may be a very complex regulation of the contractile state (by variations in inotropy), preload (by variations in venous return - venoconstriction; giving load dependent increase in contraction) and afterload by varying arterial tone (which again is often balanced by inotropy). The final variable is the cardiac output.
This means that heart rate is in itself a determinant for especially shortening rate, and more indirectly, shortening, and secondly to stroke work and cardiac output.
The HUNT3 (16) and 4 (249) are two of the largest normal single center studies of Echocardiography in the world. Both are similar in size (HUNT3 1266 vs HUNT4 1412), and with normal age distribution.
HUNT 3 | HUNT 4 | |||
Women | Men | Women | Men | |
Number | 673 | 623 | 788 | 624 |
Age (years) | 47.8 (13.6) | 50.6 (13.7) | 57.2 (12.4) | 57.8 12.4) |
BMI (Kg/m2) | 25.8 (4.1) | 26.5 (3.4) | 25 (4) | 26 (3) |
BP (mmHg) | 127 / 71 (17/10) | 133 / 77 (14/10) | 127 / 72 (18/9) | 131 / 78 (17/10) |
There was a slight difference in mean age. As many of the measurements are age related, the age distribution is important in comparing populations.
HUNT 3 was acquired in 2006 - 2008, HUNT 4 in 2017 - 2019. Thus the echo populations are two different cohorts, although there was some overlap, as participants from HUNT3 were invited to participate in HUNT4, but the individuals paticipating in both were aged 20 years, and the comparison will give further data an ageing.
Both normal studies excluded patients with heart disease, diabetes and hypertension.
HUNT3 was taken on GE Vivid 7, and analysed in EchoPAC version BT06, (except strain and strain rate, which was analysed in the proprietary segmental strain analysis software.
HUNT4 was acquired on GE Vivid E95 and analysed on EchoPAC version 203. Thus there were technical developmental differences as well.
The two studies differ in much of the measurement methodology, meaning that comparison is interesting from a methodological viewpoint, but also in looking at age and sex relations across methods. In linear dimension measurements, HUNT 3 used mainly M-mode, HUNT4 B-mode.
Dimensions of the ventricle is closely related to the functional measures. While the motion indices of displacement and velocity are dimension unrelated, strain and strain rate are relative deformation measures, and thus related to dimensions. Thus changes in dimensions will relate to changes in strain and strain rate. The HUNT study, being ta large study of normals has published normal values, related to age and gender (19):
Conventional left ventricular cross sectional measures from M-mode in the HUNT3 study by age and gender, raw and indexed for BSA. SD in parentheses.
Age (years) | N | IVSd | IVSd/BSA | LVIDd | LVIDD/BSA | FS (%) | LVPWd | LVPWd/BSA | RWT | RWT/BSA |
Women | ||||||||||
<40 | 207 | 7.5 (1.2) | 4.2 (0.6) | 49.3 (4.2) | 27.5 (2.6) | 36.6 (6.1) | 7.7 (1.4) | 4.3 (0.6) | 0.31 (0.05) | 0.17 (0.03) |
40–60 | 336 | 8.1 (1.3) | 4.5 (0.7) | 48.8 (4.5) | 27.3 (2.8) | 36.5 (6.9) | 8.3 (1.3) | 4.6 (0.7) | 0.33 (0.05 | 0.19 (0.03) |
> 60 | 118 | 8.9 (1.4) | 5.1 (0.8) | 47.8 (4.8) | 27.4 (3.1) | 36.0 (9.1) | 8.7 (1.4) | 5.1 (0.8) | 0.37 (0.07) | 0.22 (0.04) |
All | 661 | 8.1 (1.4) | 4.5 (0.8) | 48.8 (4.5) | 27.4 (2.8) | 36.4 (7.1) | 8.2 (1.4) | 4.6 (0.8) | 0.34 (0.06) | 0.19 (0.04) |
Men | ||||||||||
<40 | 128 | 8.8 (1.2) | 4.3 (0.6) | 53.5 (4.9) | 26.1 (2.6) | 35.5 (6.9) | 9.2 (1.3) | 4.5 (0.7) | 0.34 (0.06) | 0.17 (0.03) |
40–60 | 327 | 9.5 (1.4) | 4.6 (0.7) | 53.0 (5.5) | 26.0 (3.0) | 35.8 (7.4) | 9.7 (1.4) | 4.7 (0.7) | 0.37 (0.07) | 0.18 (0.03) |
> 60 | 150 | 10.1 (1.6) | 5.1 (0.9) | 52.1 (6.4) | 26.3 (2.9) | 36.0 (8.0) | 10.0 (1.3) | 5.1 (0.7) | 0.39 (0.07) | 0.20 (0.04) |
All | 605 | 9.5* (1.5) | 4.6† (0.8) | 52.9* (5.6) | 26.0† (2.9) | 35.8 (7.5) | 9.6* (1.4) | 4.7† (0.7) | 0.37 (0.07) | 0.18 (0.04) |
Total | 1266 | 8.7‡ (1.6) | 4.6 (0.8) | 50.8‡ (5.4) | 26.7 (2.9) | 36.1 (7.3) | 8.9 (1.6) | 4.7 (0.7) | 0.35 (0.07) | 0.18 (0.04) |
*p<0.001 compared to women. †p<0.01 compared to women. ‡Overall p<0.001 (ANOVA) for differences between age groups. RWT: relative wall thickness.
Wall thicknesses and LVIDD correlated with BSA (R from 0.41 - 0.48), Thus, all values were consistently higher in men due to this. FS, of course, did not correlate with BSA, and was thus gender independent. Wall thicknesses increased with age (R=0.33), and LVIDD decreased with age, as opposed to older studies (34, 35), possibly because of their smaller size.
FS remained constant between age groups, in accordance with other studies.
The values are M-mode derived.
In HUNT3, the upper normal cut off for RWT would be 0.49, which is relevan for M-mode derived values.
Relative wall thickness is generally considered to be a body size independent measure, as both wall thicknesses and LVIDD are body size dependent, the RWT, supposedly, is normalised for heart size, and hence, for body size. Interenstingly, in the HUNT study this was not the case, although correlation with BSA was very modest (R=0.18). This probably do not warrant normalising RWT for BSA. More pronounced was correlation with age (R=0.34), increasing from 0.31 to 0.37 in women, and from 0.34 to 0.39 in men. For M-mode derived values this means that the cut off for RWT would be 0.51 in the upper age groups, for both sexes.
The age dependency is a logical consequence of the unchanged LVIDd and increasing wall thickness, and has been shown also previously (33).
Relation of RWT and BSA in HUNT3. This shows that RWT is not perfectly aligned with body size. | RWT and age in HUNT3. This shows a more marked dependence of RWT and age, so age related normal values is probably warranted. |
Comparing with the values from HUNT4 (249), which were measured in 2D the values according to aghe and sex can be found in the original publication.:
Age (years) | IVSd (mm) | LVIDd (mm) | LVIDs (mm) | FS (%) | LVPWd (mm) | RWT |
Women | ||||||
20 - 39 | 6.8 (1.2) | 49 (4) | 33.3 (4.0) | 0.32 | 6.6 (0.9) | 0.27 |
40 - 59 | 7.4 (1.3) | 48 (4) | 33.0 (3.7) | 0.31 | 6.9 (1.0) | 0.30 |
60 - 79 | 8.1 (1.5) | 45 (4) | 30.9 (4.2) | 0.31 | 7.5 (1.1) | 0.35 |
> 79 | 8.2 (1.0) | 41 (4) | 28.1 (3.0) | 0.31 | 7.7 (1.2) | 0.39 |
All | 7.7 | 47 | 32 | 0.32 | 7.2 | 0.32 |
Men | ||||||
20 - 39 | 7.9 (1.3) | 52 (4) | 35.2 (4.3) | 0.32 | 7.3 (1.0) | 0.29 |
40 - 59 | 8.7 (1.3) | 52 (5) | 35.9 (4.5) | 0.31 | 8.0 (1.2) | 0.32 |
60 - 79 | 9.2 (1.5) | 50 (5) | 33.9 (4.8) | 0.32 | 8.3 (1.2) | 0.35 |
> 79 | 9.3 (1.7) | 48 (6) | 34.6 (3.9) | 0.28 | 8.2 (0.8) | 0.36 |
All | 8.9 | 51 | 34.9 | 0.32 | 8.1 | 0.33 |
Total | 8.2 | 49 | 33.3 | 0.31 | 7.6 | 0.33 |
Values are taken from (249). All age differences were significant) Values were corrected for the numbers in each age class before averaging by me. FS and RWT are calculated by me from the basic measurements, and likewise corrected for numbers.
What we see is that the M-mode derived values of HUNT3, are slightly higher than the 2D derived values from HUNT 4, for wall thickness and chamber diameter. The sex and age distribution of the subjects are about the same in both studies, although the age related decline in LVIDd is steeper in HUNT4.
The NORRE study (250), however, had intermediate values; IVSd 8.6 mm and LVPWd 8.8 mm, although closer to HUNT3. LVIDd was fairly close in HUNT 3 and 4, somewhar smaller in NORRE (44.3 mm), age related values for linear dimensions are not given. Mean FS in NORRE can be calculated to 33%.
Thus, the NORRE study seems to confirm the bias between M-mode and 2D derived measures, being lower than HUNT3.
This consistent with a statistical bias towards skewed measures in M-mode, which tend to be corrected in 2D (224), so values in HUNT3 may be valid for M-mode measures, but will overestimate the true values, and is not transferable to 2D derived measures.
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Reconstructed M-mode with a fairly straight cross angle between the M-mode line and the LV long axis. | Reconstructed M-mode from the same loop, but with the M-mode line crossing the LV long axis at a skewed angle, showing thicker walls and wider cavity, due to the angulation. | B-mode measurement across the ventricle in the same loop. Wall thicknesses are similar to the straight angle M-mode. LVIDd (and hence, RWT) are slightly different as the measurement line does not cross the posterior wall at exactly the same point. |
However, the relation between values, and relations with age, seems to be valid across the methods. Decreasing LVIDd but increasing WT with increasing age was found in both HUNT3 and 4, so these relations are not method specific. FS do not seem to differ very much between age groups in either study. Theoretically, skewed measurements would not affect the ratio of the measured values (RWT and FS). However, as we see, both RWT and FS are slightly lower in HUNT4, and in the NORRE study the mean FS was 0.33 as well. Skewed measures should not give increased ratios per se, but, angulation will also be affected by the AV plane motion, so there are more systematic errors than just angulation.
The cut off for RWT in HUNT 4 would be ca 0.47 over all. Mean RWT in the NORRE study can be calculated to 0.39, higher than both HUNT 3 and 4, confirming that the cut off seems to be to low. While RWT increases with age in both studies. As RWT is derived from wall thicknesses which increases significant with age (249) and LVIDd which decreases significant with increasing age, (249), the increase in RWT has to be significant as well.
Increasing wall thickness and LVIDd with higher BSA, and relative wall thickness thus also have to increase with increasing age as diameter decreases and wall thickness increases in both studies. The cut off for RWT in HUNT 4 would be ca 0.47 over all, and increasing by age, possibly around 0.50 for the upper age groups, still higher than previous normal cut off values (224).
The age relation is not taken into account in current guidelines, but as upper normal limit increases by age, age related values should be warranted.
Left ventricular length and external diameter is also important in an evaluation of the total strain images. In the HUNT3 study, we initially measured wall length, from the mitral annulus to the apex, which will over estimate the LVL, but is proportional (19). Later, we refined this to an ellipsoid model, allowing to calculate the mid cavity LV length (65) as shown below.
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Left ventricular length. Wall lengths were measured in a straight line (WL) in all six walls from the apex to the mitral ring. This wil underestimate true wall lengths (dotted, curved lines), but will be more reproducible, as the curvature may be somewhat arbitrary. LVL was calculated as mean of all four walls, thus overestmating true LVL (yellow line) slightly, but again the arbitrary placement in the middle of the ostium will result in lower reproducibility, while taking the mean of six measurements will increase it. | Ellipsoid model of the left ventricle. All basic measures are linear, and the ellipsoid model assumes symmetrical wall thickness, declining to half in the apex, mitral annular diameter constant; equal to ventricular end systolic diameter, as LV diameter decreased by 12.8% is systole while the fibrous mitral annulus may be assumed to be more constant. LVL is calculated by the pythagorean theorem, using 1/2 LVIDd plus 1/2 WTd. |
Left ventricular external diameter, is simply the sum of the wall thicknesses and LVIDd. From the geometric calculation, we found LV internal and external lengts (assuming apical wall thickness in both systole and diastole to be 50% of basal wall thickness:
Left ventricular wall length and external diameter by age and gender from the HUNT3 study, raw and indexed for BSA..
Age (years) | N | LVEDD (cm) | LVEDD/BSA (mm/m2) | LWVL (cm) | LWVL/BSA (cm/m2) | LVWL/LVEDD | LVELd (mm) | LVILd (mm) |
Women | ||||||||
<40 | 207 | 6.45 (0.48) | 35.9 (2.7) | 9.4 (1.6) | 5.23 (1.00) | 1.46 (0.26) | 91.0 (6.2) | 87.2 (6.1) |
40–60 | 336 | 6.52 (0.52) | 36.5 (3.2) | 9.1 (1,7) | 5.08 (0.95) | 1.40 (0.27) | 88.5 (6.0) | 84.3 (5.9) |
> 60 | 118 | 6.52 (0.52) | 37.7 (3.5) | 8.9 (1.3) | 5.08 (0.79) | 1.36 (0.23) | 85.0 (5.9) | 80.1 (5.9) |
All | 661 | 6.51 (0.51) | 36.5 (3.2) | 9.1 (1.6) | 5.13 (0.93) | 1.41 (0.27) | 88.7 (6.4) | 84.6 (6.4) |
Men | ||||||||
<40 | 128 | 7.16 (0.53) | 35.0 (2.9) | 10.3 (1.7) | 5.02 (0.88) | 1.44 (0.25) | 99.6 (6.4) | 95.0 (6.4) |
40–60 | 327 | 7.22 (0.58) | 35.0 (3.2) | 10.0 (1.8) | 4.84 (0.89) | 1.39 (0.26) | 97.3 (7.4) | 92.5 (7.4) |
> 60 | 150 | 7.22 (0.68) | 36.5 (3.1) | 9.5 (1.8) | 4.80 (0.97) | 4.80 (0.97) | 92.1 (7.8) | 87.1 (7.8) |
All | 605 | 7.21 (0.59) | 35.3 (3.1) | 9.9 (1.4) | 4.86 (0.91) | 1.38 (0.27) | 96.5 (7.8) | 91.7 (7.8) |
Total | 1266 | 6.84 (0.65) | 36.0 (3.2) | 9.5 (1.8) | 5.00 (0.93) | 1.40 (0.27) | 92.4 (8.1) | 88.0 (7.9) |
SD in perentheses, LVELd: Diastolic external length LVILd diastolic internal length. p < 0.001.
It is logical that LVEDd increased both with BSA (R=0.60) and modestly with age (R=0.11, the unchanged LVIDd being part of it, dilutes the effect of wall thickness) (19).
Left ventricular length, on the other hand, increased with BSA (R=0.29), but decreased with age (R = -0.12).
Left ventricular external diameter, is calculated from the published values as the sum of the wall thicknesses and LVIDd. Values were corrected for the numbers in each age class by me.
Values from HUNT4 , measured in 2D the values according to age and sex can be found in the original publication (249).: LV lengths were exported from the volumetry tracings in 2D (Dalen H personal communication), meaning they represent inner length.
Age (years) | IVSd (mm) | LVIDd (mm) | LVPWd (mm) | LVEDd (mm) | LVILd-4ch (cm) | LVILd2ch (cm) |
Women | ||||||
20 - 39 | 6.8 (1.2) | 49 (4) | 6.6 (0.9) | 62.4 | 8.5 (0.6) | 8.6 (0.8) |
40 - 59 | 7.4 (1.3) | 48 (4) | 6.9 (1.0) | 62.3 | 8.3 (0.6) | 8.4 (0.6) |
60 - 79 | 8.1 (1.5) | 45 (4) | 7.5 (1.1) | 60.6 | 7.9 (0.6) | 7.9 (0.6) |
> 79 | 8.2 (1.0) | 41 (4) | 7.7 (1.2) | 56.9 | 7.1 (0.5) | 7.3 (0.5) |
All | 7.7 | 47 | 7.2 | 61.5 | 8.1 | 8.2 |
Men | ||||||
20 - 39 | 7.9 (1.3) | 52 (4) | 7.3 (1.0) | 67.2 | 9.7 (0.6) | 9.7 (0.6) |
40 - 59 | 8.7 (1.3) | 52 (5) | 8.0 (1.2) | 68.4 | 9.2 (0.6) | 9.4 (0.7) |
60 - 79 | 9.2 (1.5) | 50 (5) | 8.3 (1.2) | 67.5 | 8.9 (0.6) | 8.9 (0.6) |
> 79 | 9.3 (1.7) | 48 (6) | 8.2 (0.8) | 65.5 | 8.5 (0.5) | 8.4 (0.5) |
All | 8.9 | 51 | 8.1 | 68.0 | 9.1 | 9.2 |
Total | 8.2 | 49 | 7.6 | 64.3 | 8.5 | 8.6 |
Values are taken from (249). All age differences were significant, although not tested for LVEDD. As in HUNT3, we se that the age effect on LVEDd is diluted, by the decreasing LVIDd and increasing WT.
In HUNT3, the mean left ventricular internal length was calculated from wall lengths to 88 mm. In HUNT4, the internal LV length was 85 mm.
Disregarding the separate age group > 80, which has no comparable group in HUNT3 (and which are small), we see that internal length decline from 87 - 80 mm in women and 95 - 87 in men, while in HUNT 4 the lengths declined from 85 - 79 mm in women and 97 - 89 in men.
Thus both studies confirm a decline in both internal diameter and length with increasing age.
Fundamental findings are summarised below:
Fundamental findings in the HUNT study: With increasing BSA, both wall thickness, internal diameter (and hence, external diameter) and relative wall thickness increase, showing that neither measure is independent of body size (or heart size). The length / external diameter, however, remains body size independent, being a true size independent measure. Differences are exaggerated for illustration purposes. | With increasing age, both wall thickness (and hence, external diameter) increase, while internal diameter is age independent. Left ventricular length decreases, and hence length / external diameter decreases, and i a measure of age dependent LV remodeling. This has implication for LV mass calculation. Dimension changes are exggerated for illustration puposes. |
The ratio L/D did not correlate with BSA in HUNT3, was near gender independent (although the difference was significant due to the high numbers), but declined somewhat more steeply with age (R = -0.17). In HUNT4, the ratio between external diameter and internal length was 1.32 in women and 1.34 in men, but declined with age.
This has some important corollaries:
Applying the linear measures to an elliptical model of the left ventrcle, allowed the estimation of LV volumes (471).
Ellipsoid model of the left ventricle. All basic measures are linear, and the ellipsoid model assumes symmetrical wall thickness, declining to half in the apex, mitral annular diameter constant; equal to ventricular end systolic diameter, as LV diameter decreased by 12.8% is systole while the fibrous mitral annulus may be assumed to be more constant.
The ellipsoid model has some limitations. Being symmetric, it do not conform totally to the shape of the LV, which is assymmetric, as in other model studies.
An indication of this was that while all linear measurements were near normally distributed, there was a greater skewness in the calculated volunes:
Comparing skewnesses of the distributions of the linear measures (which is small), with the calculated volumes (which is significantly (greater), seems to indicate a systematic error in the volume data from the model.
Despite this, findings were interesting.
Age | LVEDV(ml) | SV(ml) | EF(%) | Myocardial volume d (ml) |
Women | ||||
<40 | 111.6(21.6) | 76.3(16.4) | 68(6) | 87.0 (19) |
40-60 | 106.9(21.7) | 72.7(17.0) | 68(6) | 92.8 (19.6) |
>60 | 97.9(19.7) | 65.4(16.9) | 66(9) | 95.6 (18.9) |
Total | 106.8(21.8) | 72.6(17.3) | 68(6) | 91.4 (19.6) |
Men | ||||
<40 | 144.8(30.5) | 96.1(22.9) | 66(8) | 125.3 (23.6) |
40-60 | 138.1(31.1) | 92.2(23.8) | 67(8) | 129.7 (25.3) |
>60 | 126.3(33.7) | 84.1(25.7) | 66(8) | 128.2 (26.8) |
Total | 136.6(32.2) | 91.0(24.4) | 67(8) | 128.4 (25.3) |
All | 121.1(31.1) | 81.4(22.9) | 67(8) | 101.4 (27.9) |
LV volumes in HUNT4
Volumes in HUNT 4 are 2D measurement derived, values given are taken from (249).
Comparing with the values from HUNT4 (249), which were measured in 2D, the values according to age and sex can be found in the original publication.:
Age (years) | LVEDV(ml) | SV(ml) | EF(%) |
Women | |||
20 - 39 | 114 (26) | 68 | 60 |
40 - 59 | 102 (19) | 62 | 61 |
60 - 79 | 84 (19) | 51 | 60 |
> 79 | 67 (7) | 41 | 62 |
All | 94 | 57 | 61 |
Men | |||
20 - 39 | 145 (28) | 84 | 58 |
40 - 59 | 136 (29) | 81 | 60 |
60 - 79 | 119 (27) | 71 | 60 |
> 79 | 104 (18) | 62 | 59 |
All | 128 | 76 | 59 |
Total | 109 | 66 | 60 |
SV is calculated from EDV and ESV, corrected for the numbers in each age class before averaging.
Compared to the ellipsoid model from M-mode derived values, LVEDV is lower, but the relation to age, with declining values in increasing age are the same. This of course follows from the simultaneous age dependent decrease in both LVIDd and LVILd by age in both studies. Myocardial volume is not calculated in HUNT4.
The NORRE study shows even lower EDV, but LVEDV decrease with age too.
As we have already shown, left ventricular wall thickness increased with age, LV diameter was unchanged, while LV wall length decreased (19). However, LV volume increased by age (65), this effect being more profound in women.
Myocardial volumes were not calculated in HUNT4. The NORRE study gives increasing myocardial mass (which in practice is only myocardial volume × 1.05) in women, with mean 146 g in men and 112 g in women.
But tThe HUNT 3 population despite exclusion of patients with history or treatment for hypertension, had an increasing mean SBP and DBP with increasing age, due to an increasing number with BP above hypertensive levels:
A: Mean BP showing an age related increase, above 60 about half is in the hypertensive level >140/90. B: LV volume in the different BP groups (results were not different if 140/90 was used). There is significant higher volumes in the >130/80 group, but in neither group was there any significant increase with increasing age.
There was a weak, but significant correlation of LV volume with age (R=0.14, p<0.001), but neither in linear regression nor partial coorrelation was there any significant increase with age, indicating that the age effect is mainly an BP effect.
Myocardial compressibility in relation to strains is discussed in the fundamental concepts section.
The volume ratio by strains is
Given myocardial incompressibility,
,
if the myocardium is compressed during systole,
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In the HUNT3 study, using the strain product on linear measures, the strain product, being equal to the volume ratio was 1.009 (1.0136 - 0.99851) using straight line wall measures (longitudinal strain -16.3%), and 0.9957 (1.003 – 0.98896) using mid ventricular line (longitudinal strain -17.1%). However, speckle tracking tends to measure higher GLS, because of the shortening due to inward tracking of the wall thickening, and wall thickening varies too much between studies to give any meaning of the strain product at all. The answer cannot be given by strains. For speckle tracking, we know that the resolution, and hence the tracking is different in the axial and lateral direction, so the values are not necessarily inter related in a proper way,
In the model study, however, myocardial volumes could be estimated. Here, we found a myocardial volume reduction in systole of 3.28 ml, or 2.5% of myocardial volume, 4.8% of SV.
This corresponds to a Vs/Vd of 0.975 (SD 0.112), 95%CI ((0.969-0.981) .
But as the model has limited accuracy, this is not normative either. Our main finding was that this compressibility, however, was not related to age, BP or BSA.
The full heart cycle is composed of the interactions between tissue deformation, flow and pressure in a complex manner. The intraventricular flow is an important part of the total picture.
Combined high framerate tissue Doppler and vector flow imaging, showing a vortex in the LV. Image courtesy of Annichen S Daae.
The main intervals of the heart cycle is defined by the valve closures and openings. That means start and end of flow, but with the right positioning of the sample volume /beam, it can also capture the valve closure clicks.
With the right positioning if the sample volume, all left sided valve closures and openings can be registered simultanepously. Valve closures can be seen as clicks, and valve openings can be seen by the start of flow.
Valve opening and closures then can define the main intervals of the heart cycle, IVC, LVET, DFP and IVR.
Pulsed wave Doppler with sample volume positioned between the aortic and mitral ostia. Both start and stop of the flows, as well as the valve closure clicks are seen, dividing the heart cycle into the four main phases, ejection, diastolic filling, and between them the isovolumic contraction and relaxation phases..
Tissue Doppler and deformation imaging, howebver have given increased understanding of the basic physiology. Looking at the short phases of the heart cycle, like pre ejection, isovolumic relaxation, and early and late diastole, velocity and strain rate are the most useful, as explained here.
The pre ejection period, is the period from start of ECG, to the AVO (i.e.) the start of ejection (110). It consists of:
The electromechanical activation consists of electrical conduction of the signal from the AV-node through the His' bundle and the anterior and posterior left hemi-bundles. Early experimental and invasive studies seemed to show that there is initial endocardial activation almost simultaneously in mid septum and mid lateral wall, after 0 - 15 ms after onset of ECG (111, 112), but this will be partly concomitant with electromechanical delay at the cellular level. Then, there is electromechanical delay at the cellular level, the action potential generating Calcium influx, again generating release of more calcium form the SR, resulting in onset of cell contraction through actin-myoisin cross bridges. This process takes about 20 - 30 ms (113).
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Excitation-tension diagram. After Cordeiro. The Action potential triggers the influx of calcium, which triggers further release of Ca2+from sarcoplasmatic reticulum. Calcium binds to troponin, and allows activated (by ATP) myosin heads to bind to troponin sites on actin (cross bridge forming) and release energy, causing the filaments to slide along each other, as long as there is a high calcium concentration in the cytoplasm. | Image of beating isolated myocyte. The myocyte is treated with an agent that fluoresces in the presence of free calcium in the cytosol. We see that the cell lightens and shortens simultaneously; stimulation causes an increase in free calcium (released mainly from the sarcoplasmatic reticulum), causing the cell to become lighter. The free calcium is the trigger for the binding of ATP, and the formation of activated cross bridges between actin and myosin, and the subsequent rotation and release, which leads to the buildup of tension in, or shortening of the cell. Image courtesy of Ph.D. Tomas Stølen, cardiac exercise research group (CERG), |
With tissue Doppler, it became evident that before ejection, a short positive velocity spike was visible. It was visible in both septum and the lateral wall (72), and it corresponded to a very small pre ejection apical displacement as seen by M-mode and displacement traces.
Pre ejection velocity is seen as a small, positive velocity spike before the main motion of the ejection phase, and, correspondingly, a small apiucal motion that can be discerned both in the M.mode and displacement traces. This pre ejectiopn motion is present both in the septum and the lateral wall.
Pre ejection spike can be seen to be lower than peak ejection in most instances | ||
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-by spectral tissue Doppler | by colour tissue Doppler | and by speckle tracking. |
The timing of this event has been shown to precede MVC (72).
The pre ejection spike thus occurs BEFORE MVC, as seen here (valve openings and closures from Doppler flow - different cycles). | MVC is concomitant with the stop of pre ejection apical motion. |
It has been suggested to be passive rebound after atrial contraction. This has been suggested, but the pre ejection spike is seen also in atrial fibrillation (114, 117). Rebound after atrial relaxation is present, though, and can be demonstrated, both with and without (as in AV-block) succeeding ventricular contraction, but the timing in relation to the following systole depends on the PQ time. What we see, is that after the a' wave, there may be some rebound oscillations visible in long PQ time or dropped beats, which normally would be dampened by the stiffening of the ventricular myocardium by the onset of contraction.
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Ultra high frame rate tissue Doppler from the base of the septum a subject with atrial fibrillation. Even with no atrial activity, there is pre ejection velocity spikes, showing them to be ventricular in origin. | Ultra high frame rate tissue Doppler from the base of the septum a subject with 1st degree AV block. (This is a highly trained, healthy subject, the AV block is physiological)Three spikes are seen before ejection (arrows). Here, the initial spike must be atrial recoil, coming before start of the the QRS, it cannot be ventricular i origin. Even the second spike may be atrial, or a fusion of an atrial bounce and ventricular contraction. | Ultra high frame rate tissue Doppler from the base of the septum a subject with 2nd degree AV block, as seen by the second P-wave following the first heartbeat, with no QRS nor ejection velocities. The atrial recoil can be seen as three velocity spikes (arrows), indicating that the mitral ring bounces. However, this is in a situation without LV myocardial tension. At start of the first heart cycle, there may be some fusion between atrial recoil and ventricular contraction as seen by the timing. |
If PQ time is very long, the atrial rebound and pre ejection spikes become separated:
PW Doppler of mitral flow and tissue Doppler from a patient with a PQ time > 400 ms. This pushes the A wave forward to the preceding heartbeat, causing EA fusion as explained elsewhere, but increases the interval between atrial and ventricular systole. After the a' wave in tissure Doppler there is a clear positive wave, that can be seen before the normal pre ejection spike, and may represent a rebound from atrial contraction. With HIS pacing on, the PQ-time is normalised, and both EA fusion and the rebound wave are abolished, the latter may be dampening with the onset of ventricular myocardial contraction (stiffening). The normal pre ejection spike, however, is enhanced. Images courtesy of Prof. Dr. med. H. Kühl, Chefarzt, Klinik für Kardiologie und Internistische Intensivmedizin, Klinikum Harlaching, München.
With ultra high frame rate (1200 FPS) ultrasound, (HFR IQ data), where both MV M-mode and TDI was reconstructed from the same cycle, and the velocity spikes were comp+ared to a reconstructed MV M-mode, MVC was seen to come after the pre ejection spike. In a study of ten healthy subjects, time intervals from start of ECG to start of the initial pre ejection velocity spike in the septum was 22.7 ms, and from this to MVC was 29.6 ms. Findings were very consistent with the electrophysiologic intervals, indicating that the pre ejection spike was active contraction, but also that the onset of active contraction occurred before MVC (114).
Ultra high frame rate tissue Doppler (about 1200 FPS) from the base of the septum a normal subject. The timing is evident, with ECG starting first, then the pre ejection velocity spike starting about 23 ms later, and then the mitral valve closure about 30 ms after this. This recording is from the septum, and as can be seen, in the septum there is a second spike before ejection starts. It can be seen to repeat from beat to beat. This was not present in the lateral wall.
The findings were later confirmed by a simple study with transfer of valve motions from Doppler flow recordings to tissue Doppler (72). This method is slightly less accurate, both because of the lower frame rate, and as there is some heart rate variability when valve clicks are transferred from other cycles.
Valve closures by valve clicks, and openings by onset of Doppler flow, to the right transferred to the analysis window for tissue Doppler to the left. Relations between timing of tissue Doppler motions and valve motions are typical: Q to onset pre ejection spike (EMD) was about 25 ms, duration of the pre ejection spike about 50 ms, and MVC to end pre ejection spike (ms) about 10 ms.
However the findings were close; Pre ejection spike:
Thus, pre ejection spike is not isovolumic contraction. This rather absurd explanation is a contradiction in terms, as it is present both in the septum and lateral wall, it must necessarily correspond to a volume reduction, which means that the phase is not isovolumic at all. Still, there were some publications on both isovolumic velocity and isovolumic acceleration as contractility measures based on this erroneous concept, as discussed below.
The pre ejection shortening has so short duration, that it is most readily demonstrated by strain rate rather than strain.
Pre ejection strain rate, active contraction can be seen to be simultaneous in both walls and in both basal and apical level, and to occur before MVC.
In 1973 by phonocardiography (115), and in 1978 by radioopaque markers (116) the ventriculoatrial crossover was demonstrated to occur ca 40 ms before MVC, demonstrationg that this is the onset of thesion buildup, i.e. that this pre ejection motion is active contraction.
The finding of mitral ring motion in both walls, must mean that this is a real volume reduction of the LV (72), again indicating that this is active contraction. Even without mitral flow, the displacement of the mitral ostium will exclude (atrialize) a small volume from the ventricle:
The apical ring motion is equivalent to a longitudinal shortening, as shown by the strain rate above, and thus a volume reduction, not by flow, but by exclusion of a small volume by the ring motion while the mitral valve is still open.The ring motion will also displace the mitral leaflets towards the base and middle, as they move in the blood. However, the main mechanism for MVC seems to be intraventricular flow as discussed later.
This volume reduction before MVC was also demonstrated experimentally in dogs by conduction catheter, and was shown to be ca 4.2% of EDV (117).
The presence in both septum and lateral wall, and the fact that both walls show negative strain rate (72), as well as the concordance with the known electrophysiology (111 - 113), makes it also very probable that the pre ejection spike is active contraction, and shows the true electromechanical activation, by the left anterior and left posterior bundles.
The pre ejection motion, is thus a contraction, representing a shortening before MVC, meaning at a load of atrial pressure, i.e. near unloaded. Peak pre ejection (72):
Peak pre ejection | Septal | Lateral |
Velocity (cm/s) | 3.59 (1.56) | 3.55 (1.58) |
Strain rate (s-1) | -0.77 (0.61) | -0.66 (0.38) |
Standard deviations in parentheses.
Could this be a measure of contractility?
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Isotonic isometric twitches. The peak tension development is slightly after onset of contraction, | Pre ejection velocity spike occurs before MVC, and thus occurs at the level of atrial pressure, as close to an unloaded situation as one gets. |
Although some papers have been published with "isovolumic acceleration" as contractility measure, it isn't.
Logically, the MVC should be at the abrupt stop of the pre ejection displacement, i.e. where the velocity trace crosses the zero line after the spike due to the sudden increase in resistance to movement at the closure ov the valve as the cusps stay in the stationary blood stream :
Blown up image of the pre ejection period, displacement to the left, velocity to the right. The logical time for the MVC is when the apical displacement stops abruptly. This is equivalent with the velocity spike retuning to zero, as can be seen with the relation to ECG. As we see, there is a slight initial velocity, and then a break point in the velocity curve, concomitant with the onset of shortening as seen by the displacement curve, corrsponding to the second spike as seen below.
Thus it's the MVC that is actually terminating the pre ejection longitudinal motion /volume reduction.
An experimental study, where the MV was stented open, confirmed this, showing that while shortening started at the same time in the stented situation, the motion (and strain) continued directly into the ejection phase, without the abrupt stop after pre ejection velocity, showing that the start of pre ejection shortening is the true electromechanic activation, while it is the MVC itself that terminates the pre ejection spike (117):
Representation of the findings in the experimental study, showing that as onset of contraction results in shortening, this is interrupted by MVC (onset of IVC), and the velocity spike returns to zero. When mitral valve is stented, this results in a smooth non-interrupted transition to ejection velocities, so the pre ejection strain and velocity (red part of curves) are nearly obliterated, and the onset of the ejection phase the major volume rediction seen by the major downward stroke of the strain curve and the upward stroke of the velocity curve) is starting at the point of the previous MVC, thus the ejection phase now starts at the onset of what was IVC without the stent..
Thus, the pre ejection velocity or strain rate spike is not in accordance with the physiological expectations, and peak protosystolic velocity is not a close measure of maximum unloaded shortening velocity / rate.
The reasons for this may be that as an event of very short duration;
The main point is that the peak protosystolic velocity or acceleration is not a contractility measure.
It has been suggested that the pre ejection annular motion is the mechanism for MV closure (117), the apical motion of the annulus on the stationary blood forcing the valves in the basal direction. However, the basal mitral leaflet motion is greater than the apical motion of the annulus, so this is insufficient to explain the closure. In addition, this hypothesis doesn't take the intraventricular flow into consideration.
During pre ejection, concomitant with the QRS, there is an intraventricular vortex that stems from the early and late filling vortices.
At the end of diastolic filling, the flow in the ventricle consists of a vortex formed by merger of the vortices from the early and late filling phases as discussed later. This vortex is counterclockwise seen in the traditional 4-chamber view, with basally directed flow along the septum, and apically directed flow along the lateral wall (118, 119).
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Septal colour M-mode showing basally directed flow along the septum during PEP. It can be seen to start at the beginning of MV closure, and is directed towards the anterior mitral leaflet. | Vector flow imaging, showing the intraventricular counterclockwise vortex during pre ejection, already before MVC. The finding is consistent with the colour M-mode findings. Image courtesy of Annichen S Daae. | Lateral colour M-mode showing apically directed flow along the lateral wall during PEP. |
The vortex is the end result of the diastolic filling (118), as discussed later. The basally directed flow along the septum will contribute to two mechanisms:
The true isovolumic contraction time (IVC) is defined from MVC to AVO, and MVC is defined by the true valve closure, while the AVO is marked by the start of LVOT flow:
As shown above, MVC is at the end of pre ejection tissue velocity:
Thus, in this phase there is no volume change, and, hence, no deformation. This phase it on the other hand, the period of most rapid pressure rise, peak dP/dt, which occurs during IVC, close to the AVC (121). This represents the most rapid rate of force development (RFD). Peak dP/dt can be measured from flow velocity if there is a small MR (122).
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Peak rate of pressure rise, which is the closes correlate to the rate of force/tension development. This occurs during IVC | dP/dt can be measured by the velocity increase, if there is a small MR, too small for generating a pressureincrease in the atrium. It is customary measured between 1 and 3 m/s, which is equivalent to a pressure increase of 32 mmHg, and the dP/dt becomes a function of the time interval between then, and is used as proxy for peak dP/dt |
As it occurs before AVO, it is not afterload dependent (121), and is a useful invasive index of contractility. However, as seen fom the length force relation above, this maximal force measure is not preload independent.
During pre ejection, the vortex is seen to persist after MVC, and the septal part aligns with left ventricular outflow (118). This adds momentum and kinetic energy to the ejection flow.
Septal colour M-mode showing basally directed flow along the septum during PEP. It continues until AVO, adding momentum to the ejection, while the lateral, apically directed part of the vortex seem to attenuate.
This means there is a momentum towards the base, aligned with the LVOT even before the AVO. The velocity momentum is equivalent to a partial pressure gradient, and thus contributes to a small reduction in afterload, facilitating the ejection, by conservation of kinetic energy from filling in the vortex.
Vorticity is a measure of the rotation of the blood around each point in the image at one timepoint in the cardiac cycle and is a measure of the complexity of the blood flow. The unit of vorticity is Hz. The time trace of vorticity is found by averaging the region of interest (the LV). In our application, this is calculated by the curl or momentum of
the blood velocity field.by the formula:
As it is measured over a 2D area, it may differ from volume based calculations. The main interest is in the changes in vorticity during the heart cycle, and the relation to the kinetic energy.
Vorticity plot, showing that vorticity peaks during pre ejection and declines during ejection, despite the ejection showing higher kinetic energy. A reasonable assumption is that kinetic energy is transferred from the vortex to kinetic energy in ejection flow. Images courtesy of Morten S Wigen.
Vorticity is a quantitative measure. Mean vorticity was found in our study to be 10.0 Hz (9.2 - 11.1) in diastole and 8.6 (7.8 - 10.1) in systole (118). Again vorticity will probably differ between applications. The qualitative description can be seen from the 2D vector flow above, reasoned from the colur flow M-mode above, and even seen from the pulsed wave Doppler flow:
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The vortex is present in diastasis, but increases during late filling. Image courtesy of Annichen S Daae | pw Doppler, showing that both early and late inflow are diverted into the LVOT by a slight delay, which visualises the vortices related to the inflow. | This diversion is related to the basal motion of the AV-plane, as seen by the colour M-mode. |
Thus, qualitative clues to the vorticity are present. The presence of vortices in large parts of the heaqrt cucle, is assumed to conserve kinetic energy form one phase, to be utilised in the next, as discussed under the different phases.
It is evident that the kinetic energy is closely related to flow velocity.
The energy unit is J/m, as we do an integration of 2 dimensions only. As we see, there is a close correspondance between the velocity curves and the energy traces.
The kinetic energy per volume of blood is given by the formula:
where is the density of blood and v is the flow velocity. Blood speckle tracking allows for estimating the total energy, not only the velocity components along the ultrasound beam as in Doppler. But we do not have full volumetric data, we integrate over a 2-D region (the left ventricle), with the resulting unit of J/m.
The mean energy is found in our study was 0.21 (0.18 - 0.25) J/m in systole and 0.20 (0.17 - 0.23) J/m (118). The values will differ between applications, as the spatial algorithms will vary, and the unit J/m, which is the 2D measure differs from the true volume measure.
When inflowing blood is diverted into a vortex, the kinetic energy is conserved, to the degree that velocity is maintained. The vorticity is thus a measure of energy conservation. Peak vorticity can be seen to coincide with, but slightly later than the peak vortex inflow in LVOT:
There is some kinetic energy at the onset of early filling, which is expected from the mechanics of the IVR. Vorticity at end ejection IVR is at a minimum, although not absolutely zero, some vorticity generated at end ejection is still present. As we see, there is a close correspondance between the velocity curves and the energy traces, and between the vorticity curves and the vortex inflow in LVOT.
Flow velocity depends on the pressure gradients.
Pressure gradients can be estimated from the flow velocity field:
is the density of blood, about 1060 kg/m3 , (vx, vy) is the velocity vector and
is the blood viscosity; about 0.004 Pa/s.
Intraventricular pressure gradients were calculated along a manually defined path from base to apex (118):
Of course, calculating pressure gradients from flow, doesn't explain flow pattern, as the pressure gradients are derived from the flow pattern at the outset. Thus, thie explanation would be tautological, it is just a description of flow velocity in terms of pressure.
Positive intraventricular pressure gradients (pressure increase) towards the apex will accelerate blood flow towards the base/LVOT (ejection), and decelerate blood flow towards the apex (inflow). Negative intraventricular pressure gradients (pressure decrease) towards the apex will accelerate blood flow towards the apex (inflow) and decelerate blood flow towards the base/LVOT (ejection). There is a correspondence between velocity, as velocity increases, pressure falls, as velocity decreases, pressure increases.
Pressure gradients derived from flow velocities, top mean values, bottom all individual traces. The positive gradient means pressure increases from base to apex/decreases from apex to base, the negative gradient the opposite; pressure increase from apex to base/decrease from base to apex. This means a positive gradient is a driving force for flow from apex towards the base (LV outflow), and a decelerating force for flow from the base towards the apex (LV inflow), a negative gradient is a driving force for flow from the base towards the apex (LV inflow) and a decelerating force for from apex towards the base (LV outflow). Ejection has a biphasic pattern with positive - negative gradient, both early and late filling has the opposite biphasic patternnegative - positive gradient. Images courtesy of Solveig Fadnes and Morten S Wigen
Thus, there is also a dynamic interplay between potential energy (pressure) and kinetic enery (flow velocity) in addition to the interplay between vortex energy and laminar flow energy throughout the heart cycle as well.
The relation between flow acceleration and deceleration and pressure is also discussed below.
With high frame rate, we found that the pre ejection spike was double in the septum, but biphasic in the lateral wall. The second spike, however, was after the MVC, so this is not the atrial rebound and then the pre ejection contraction, tyhe first spike is bilateral and represents the pre ejection contraction, the second being in IVC. This is also evident as the double septal spike is seen even in atrial fibrillation.
Simultaneous recirdings with UHFR TDI from the septal (top) and lateral (bottom) base. Each recording consists of two velocity curves from two points along the Rx beam, with a distance of 1 cm. Green is the most basal. Thus, the offset between the curves represent the strain rate. The second spike seen in the septum, is not present in the lateral wall, where there is a negative velocity instead. The secons spike is after MVC, i.e. in the IVR. As this motion is reciprocal in the two walls, it represents a rocking motion, not a true volume reduction..
Ultra high frame rate tissue Doppler from the base of the septum a subject with atrial fibrillation. Even with no atrial activity, there is the same pattern of double velocity spikes, showing both to be ventricular in origin.
Tissue Doppler from basal septum of healthy subject. at 93 FPS, only one pre ejection spike can be seen, due to undersampling. By using narrow sector and maximum frame rate, it is possible to achieve 250 FPS, and here the pre ejection spike is evident. (That this is not a phenomenon of random noise, as the double spike is reproducible at the start of the next cycle.
Thus, the double spike is not the atrial rebound followed by the electromechanical activation, by the electromechanical activation followed by the IVC rocking motion. The physiological significance of this, however, is uncertain.
The apex beat can be felt through the chest wall by clinical examination. This, in itself, shows that the apex moves towards the chest wall in systole, although only a minimal amount.
The motion was first demonstrated by apexcardiography (10)
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The apex beat, shown here in an apexcardiogram recorded with a pressure transducer, demonstrating that the beat is a systolic event. (Image modified from Hurst: The Heart). | But the B-mode and M-mode echoes shows the apical motion towards the chest wall to be minimal | |
The apex beat can also be demonstrated by M-mode and tissue Doppler:
Apex beat from two different healthy subjects. Left: M-mode, middle apical displacements (from integrated velocities), and right: apical velocities.
The apex beat is seen from the pre ejection period, through the ejection. It seems logical that the apex beat is caused by the rereaction from the ejection, ocurring in the opposite direction. However, the onset of the apex beat precedes ejection.
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Resolving the motion, we see that the anterior motion in this case starts even before the start of QRS (A). (The motion seen in the displacement curve starts below zero because the tracking is set at zero by the ECG marker). Peak forward velocity (B1) is just after the QRS, while the motion stops in systole (B2), but the apex remains in the anterior position. At end systole (by T-wave in ECG), there is the start of backward motion (C), and the apex returns to the diastolic position at D. | Comparing the apex tissue velocity with LVOT flow (aligned by ECG), both start and peak apical velocity occurs before start ejection, but continues into ejection. In this case, even a second peak may be seen starting at start ejection, indicating a second impetus from ejection recoil. | And for illustration the relation between apical displacement and ejection. Apical displacement starts before ejection, and then continues into ejection, and maximal apical displacement is close to peak ejection velocity. The apex remains pressed to the chest wall during most of ejection, until the flow velocity is so low as to not generate sufficient recoil pressure, while the full return to diastolic position is somewhat later. |
And actually coincides with the pre ejection velocity.
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Comparing this with basal velocities, we see that the anterior motion of the apex starts in the pre ejection phase. The basal pre ejection spike, however, reaches peak before the apex velocity, which peaks close to the time of start ejection (B), by the tissue Doppler curves. Backward apical motion starts a little before end ejection (C), while end of backward apical motion is well within the early filling phase (D) | Relation between apical and basal displacement shows the same. |
while there is initial pre ejection contraction in the base and midwall, resulting the MVC as discussed above, this would tend to pull the apex away from the chest wall as there is no recoil force at that point. However, there is an impact from the blood coming nto the ventricle, which pushes the apex towards the chest wall already before the ejection as seen by the tissue Doppler. In stretch, this should mean that there has to be initial stretch of the apex during the pre ejectioon, simultaneous with shortening of the midwall and base.During early ejection there is thus an acceleration of the ventricle due to recoil from ejection.
The momentum of ejection is the mass of 2x SV (From the RV and LV), x mean outflow velocity. The opposite momentum is the same mass (Outer AV-plane area x AV-plane motion), but the velocity of the AC-plane is ony ca 1/10th of the blood velocity, so there is a surplur momentum pushing the nentricles towards the chest wall.
The first temporal derivatives of these parameters are
The spatial derivative of flow is
There is not a direct correspondence between contraction (tension) and shortening, nor between relaxation (tension devolution) and fibre lengthening.
As discussed above, all measures of systolic deformation is the result of fibre shortening, and fibre shortening is the result of tension versus load, and thus are not load independent. But the contraction tension is somewhat related to all measures of ejection.
As discussed above, the onset of electromechanical activation is the pre ejection.
The ejection phase, is the phase where blood is ejected, and thus the ventricular volume is decreased. The flow is the rate aof volume decrease, and thus the ventricular volume decrease equals the ejected volume.
As discussed above, electromechanical activation is marked by the pre ejection velocity spike, even before MVC, and long before AVO. During pre ejection, there is isometric contraction leading to tension increase, and thus to pressure increase.
Onset of the major shortening occurs with AVO (72, 121), and is thus a marker of onset of ejection, meaning the onset of the major volume reduction event during systole, but not of electromechanical activation. The onset of rapid upstroke velocity by colour tissue Doppler was found to be 3.8 (10.2) ms after onset of flow in the septum, and 1.0 (17.3) ms in the lateral wall (72). Thus, this marks AVO, and not electromechanical activation.Onset of rapid motion towards the apex, or rapid negative strain rate, are both markers of the onset of the rapid volume decrease, aka the ejection, aka AVO.
During ejection, both longitudinal shortening (MAPSE, peak S', GLS and global strain rate) are related to decrease in ventricular volume, and thus to fibre shortening. Flow is the rate of volume, and thus is another way of expressing volume change, and thus also a measure of fibre shortening. Thus, all measures are related to stroke volume, and to fibre shortening, and all are afterload dependent.
Ejection starts with AVO, which starts the volume decrease, and the flow in LVOT. Flow is the volume rate: volume per time Q = V/t and is usually given in litres per minute. This means that the flow also is the rate of volume change, during ejection thus the volume decrease rate. Even though the traditional wiggers diagrams show peak rate of volume decrease at AVO, this is not the case, as flow studies have shown peak flow (aka peak rate of volume decrease) to occur later, but before peak pressure (104). This has also been confirmed by direct volumetry (263).
Diagram showing the pressure-volume diagram together with the LVOT flow curve as above, showing the early peak flow, which coincides with the steepest part of the volume curve, )as expected when flow is the rate of volume decrease) occurring before the peak pressure.
It is a myth that blood always flows "downhill" down a pressure gradient, meaning from higher to lower pressure.
The truth is that flow is accelerated down a pressure gradient. This means that there is flow along a positive gradient as long at it's being accelerated, but there may be flow against a positive gradient as long as there is a positive velocity that is being decelerated.
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The pressure gradients are the accelerating force for blood. With a positive gradient, blood is accelerated, with a negative gradient, flowing blood is decelerated. Acceleration is the temporal derivative of velocity, velocity the temporal integral of acceleration. | Both during ejection and during early and late filling is the pattern positive-negative gradient, with peak velocity at pressure crossover. |
Acceleration means converting potential (pressure) energy to kinetic energy, while deceleration means pressure recovery.
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The curves show that only during flow acceleration to peak flow velocity is LV pressure actually higher than the aortic pressure (positive gradient), during the rest of ejection is the pressure gradient negative. However both LV and aortic pressures continue to increase, as a function of continuing contraction generating increasin myocyte tension. This work, however, generates increasing aortic tension, maintaining the negative gradient. | Flow velocity curves are consistent with this., showing a short acceleration phase, a longer deceleration phase. |
During ejection, there is a positive pressure gradient out of the LV only until peak velocity, after that, meaning almost the whole ejection phase the pressure gradient is negative, and the flow velocity is seen to be decelerating. (104, 267, 268)
Doppler gives flow as a velocity. This means the velocity shows how rapidly the blood volume moves along the path, the distance per time v = d/t and is given in m/s. The velocity says nothing about the amount of blood, however. Flow is the volume rate: volume per time Q = V/t and is usually given in litres per minute. But given constant orifice area, any changes in velocity will be proportional to the change in blood volume flow, , showing that there is a fundamental relation between cross sectional area of the flow, and the velocity:
Relation between velocity and flow with constant flow velocity. In this case, the velocity is distance / time v = d / t. During this time interval the distance times cross sectional area defines a volume V = A × d, which is the volume circumscribed by the motion with the velocity v. Flow (volume rate, volume per time) during the same interval t is Q = V / t, so Q = A × d / t = A × v.
This, however, holds only for constant flow velocity. Blood flow is pulsatile, but the fundamental equations of motion still hold:
Relation between velocity and flow with constant flow velocity. In this case, the velocity is distance / time v = d / t. During this time interval the distance times cross sectional area defines a volume V = A × d, which is the volume circumscribed by the motion with the velocity v. Flow (volume rate, volume per time) during the same interval t is Q = V / t, so Q = A × d / t = A × v.
Thus, flow = velocity × cross sectional area. In variable velocity, like the pulsatile flow in circulation, the flow equals the velocity integral times the cross sectional area. The integral can be obtained by the area under the velocity curve.
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Tracing the flow velocity curve by pulsed Doppler in LVOT through one heartbeat, gives the velocity time integral by the area under the curve. The LVOT diameter, can be measured in the B-mode. | The velocity time integral is the distance d, that somethin moving with the velocity of the traced curve moves, the stroke distance. The area A of the LVOT (assuming a circular cross section) is given by the measured LVOTdiameter. Thus, the volume of the cylinder given by d × A, equals the stroke volume. |
As long the area is constant, flow velocity and flow are proportional, and simultaneous. This means that flow velocity will reflect the flow. If area changes, there will be a discrepancy between flow and flow velocity, as flow remains constant, while flow velocity increases.
In a continuous channel, flow is contiguous, and must be the same across every cross section, i.e. independent of cross sectional area:
As the area A1 is larger than A2, in order to push the same amount of blood through A2, the velocity v2 must be higher than v1. As the flow is the same, and given by A×v for continuous, and A×VTI for pulsatile flow, the ratio of velocities / velocity time integrals is the inverse of the ratio of areas. This is the continuity equation. | Using the continuity equation, as the LVOT diameter (and area) is known, tracing the VTI of the LVOT flow (pw Doppler to do it in the correct level) as well as the VTI through the valve (cw Doppler). The VTI equals the stroke length, and the stroke length times the atra, equals the stroke volume. As the stroke volume is constant, the two cylinders have equal volume, and thus, the valve stenosis area (AVA) can be calculated by AVA = LVOT area × VTILVOT / VTIAO |
This demonstrates that as contiguous flow needs to be constant, flow velocity will vary with the cross sectional area, and thus will not be representative of flow, unless area is taken into account.
The continuity equation for aortic valve area has been validated against the Gorlin formula (233):
Fundamentally, both velocity and pressure represents energy. The potential energy in a fluid under pressure, is given by E = P × V, while the kinetic energy is E = ½ m v2. But this means that when velocity increases, this kinetic energy has to be recruited from somewhere, which is the pressure energy. Thus, as velocity increases, pressure has to drop:
As velocity increases for the same volume that passes point 1, also must pass point 2, the increase in kinetic energy pressure is recruited from the pressure energy. Thus, there is a pressure drop from 1 to 2. The full equation for the acceleration of the fluid is given in the Bernoully equation. However, it has been shown that both the flow acceleration part and the friction parts are so much smaller than the first part (which is the basic energy difference), and may be ignored. Thus, the modified Bernoully equation relates pressure differences to the square of velocity differences. And if v2 is much smaller than v1, the modified equation may be simplified even more. And of course, the simplest form, is still an acceptable approximation if P2 is much lower than P1.
The simplified Bernoully equation has been shown to be valid for pressure gradients across mitral stenosis (234, 235), aortic stenosis (236), and estimation of RV pressure from tricuspid regurgitation (237).
As discussed above, as flow converges towards a stenosis, there is little friction, meaning that there is laminar flow towards the stenosis. When flow has passed the stenosis, and into a receiving chamber where there is larger area, the velocity will decrease again. However, this may have different consequences. If there is perfectly laminar flow after the stenosis, the friction element is still so small, that the kinetic energy reverts to pressure: there is pressure recovery, and the pressure rises to the pre stenotic level. Doppler measurement will still measure the maximum pressure drop, and so will a manometer placed at the narrow part, but manometers before and after the stenosis will not register pressure drop.
On the other hand, if the flow is not perfectly laminar, there will be turbulence after the stenosis, resulting in frictional energy loss, and the velocity will decrease without restoring pressure energy, i.e. the energy is lost, and there is not pressure increase after the stenosis. In that case, there the Doppler gradient may be a true measure of pressure drop.
Full, partial and zero pressure recovery. The pressure drop corresponding to the velocity increase, is Pmax, the maximum pressure drop, given by P1 - P2. The net pressure drop through the stenosis, Pnet, however, is given by P1 - P3. Cw Doppler measures Pmax, (and so will a manometer placed directly into the stenosis, and may thus over estimate the effect of the stenosis.
Left is perfectly laminar flow through the stenosis. In this case, the post stenotic velocity decelerates without energy loss, and the kinetic energy is converted back into pressure again. Here, there is no net pressure drop through the stenosis. Driving pressure at P1 does not have to be increased to maintain pressure at P3, and the pressure drop at P2 is temporary. It must be remarked, however, that pressure recovery cannot be more than the initial pressure. | In the middle is partial pressure recovery. Some of the pressure energy converted to kinetic energy through the stenosis is lost when the flow velocity decelerates after the stenosis, in the form of turbulence resulting in friction. But some of the energy is recovered to pressure energy again. Thus, there is a net gradient over the stenosis, but this is less than the maximum gradient. The maximum gradient by Doppler will over estimate the net gradient. | Right, there is total energy loss through the stenosis, all kinetic energy due to the velocity increase through the stenosis is lost in turbulence and friction. Thus, |
As discussed above, if pressure measured across an aortastenosis is higher than normal intraventricular pressure, there has to be a net gradient, as the maximum pressure drop has to be positive.
In MR with a large regurgitation and an eccentric jet, the jet will impinge on the atrial wall, and follow the wall. The jet is deflected by the wall. The jet impinging on the wall will increase the pressure along the wall, and this pressure rise will decrease the flow velocity.
Large, eccentric MR, directed towrds the lateral wall, and then deflected along the wall all the way around the atrial roof and septum. The first red flow cvewlocity is the return of the MR, while the second is the diastolic pulmonary inflow jet.
This has been mistakenly been taken as the result of the Coanda effect, and this mistake has been perpetuated in various publications. However, the Coande aeffect is the apparently paradoxical deviation and adherence to a wall that is parallel with, or deviating away from the jet, and stay attatched to the surface, the mechanism is related to the Bernoully effect as explained below.
As volume flows into a receiving chamber, (from LA to LV, or from LV to Ao, receiving hambers are elastic, so the compliance decreases with volume received. The concepts of compiance and elastance are discussed above, and the relation to ejection below. As long as compliance and cross sectional area is constant, flow is påroportional with floa velocity, and the volume is proportional with the velocity time integral. However, an elastic receiving chamber decreases the compliance with increasing volume, meaning that for a given pressure increase, the flow volume decreases, or for a given flow volume vthe pressure has to increase.
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for a constant area and compliance, flow volume is proportional with velocity time integral, and equals the temporal derivative of volume. The volume thus equals the time integral of flow. | In an elastic receiving chamber, the compliance decreases (elastance increases) with increased filling. This means that for a given pressure increase, the flow volume decreases with increased filling, or the pressure increases to deliver the same volume. In CV disease, the compliance may decrease, and thus shifts the compliance curve downwards, the elastance curve upwards, decreasing flow volume or increasing pressure. |
Peak ejection performance can be defined as eiter:
As discussed above, the peak rate of ejection (peak flow) occurs early after AVO, and corresponds to the peak rate of volume decrease, aka the peak systolic annular velocity, aka the peak strain rate. Flow and flow velocity has an early peak, The delay after AVO is the time of the acceleration of ejection, and thus as flow iequals the rate of volume decrease, this must be slightly later than the onset of flow as well.
But as has been demonstrated, peak systolic pressure continues to rise until around mid systole (104), which means that peak flow starts to decrease while peak pressure continues to rise. To balance the pressure, peak myocardial tension has to increase as well, so the time of peak tension is the peak of systolic pressure. Why this discrepancy? This is due to the continued distension of the large vessels with the increase in total ejected volume.
So there is an initial increase in flow and rate of volume decrease. Peak flow is related to the contractility, but as shown below, it is earlier than the time of peak tension.
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Peak flow, which coincides with the steepest part of the volume curve, is the peak of volume ejection, but not the peak of LV tension. | During ejection, the volume is ejected into the aorta, which is distended, storing energy in the elastic properties. This distension requires pressure, and is part of the systolic LV work. Thus the tension (and work) continues to increase after peak flow, and the peak tension is at peak pressure. Ejection, continues however, until flow is completely decelerated, at AVC, and thus the maximal aortic distension is at end ejection. The aorta is then contracting again during diastole, acting as a diastolic pump, with energy stored from systole, using this energy as part of the whole pumping cycle. |
The timing of the peak flow, is determined by the crossover between the volume ejection, and the elastic distension of the aorta, which again is a function of both the elasticity and the peripheral resistance, and the magnitude of the peak is in part a function of the timing of the acceleration cutoff. But this also shows that the peak myocardial tension is later than the peak volume decrease/flow/flow velocity.
In fact, as flow is accelerated by a positive pressure gradient, and decelerated by a negative pressure gradient, the cut off of flow acceleration is the point of pressure crossover between the ventricle and the artery, and thus before peak pressure:
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Diagram, showing firstly, that flow is accelerated only during the first part of ejection, at pressure crossover from a positive to a negative gradient, the flow deceleration starts, before peak pressure, but the time of the crossover (acceleration cutoff) itself is a determinant of the peak flow. | As blood is ejected into the aorta, the aorta and large arteries being elastic, they are distended by the volume they receive. This volume is partly a function of the relative ejected volume (during acceleration), i.e. the relation between ejection volume and aortic (arterial) volume. The second factor is arterial compliance; C = |
However, despite the peak flow being early, and flow / flow velocity starting to decrease from peak after AVO, obviously the cumulated flow (which is the true meaning of the integral) and total ejected volume increases untill the end of ejection.
Thus, the peak flow rate is not only a function of contractility and pressure, but also of aortic volume and compliance, and not a "pure" LV measure.
Peak pressure, correspondint to peak tension, however, is obviously related to both stroke volume and aortic volume and compliance.
Peak flow velocity is simultaneous with peak flow, as the aortic orifice is considered constant during ejection. Flow velocity alone says nothing about the amount of blood, however. But given constant orifice area, any changes in velocity (m/s) will be proportional to the change in blood volume flow (l/min), showing that there is a fundamental relation between cross sectional area of the flow, and the velocity. Thus, LVOT flow will show the same early peak as flow recordings, but as shown in the HUNT study, there is not co variabiility between peak S and MAPSE.
Peak annular systolic velocity, S' and peak global SR, are both measures of the peak rate of longitudinal LV shortening, i.e. peak rate of length decrease. As most of the stroke volume is related to longitudinal shortening. they can be assumed to be closely related to peak rate of volume decrease. It can also se a close correspondence in timing of the peaks.
The image shows the early peak of flow velocity in the LVOT. By removing the wall filter and reducing the signal amplitude, the wall signals can be seen as well, and the wall velocity, being proportional with the volume rate, also shows the same early peak.
As the systolic longitudinal shortening is closely related to volume decrease, peak strain rate and peak annular velocity are likewise early events close in timing to peak flow:
When strain rate is sampled from most of the wall length, the shape is close to an inverted version of the basal velocity curve. In this case, both curves relate to the rate of volume change, and thus to the very early peak flow and the steepest parts of the annulus displacement and strain curves.
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Flow velocity of LVOT. This is closely related to flow, showing an early peak during the ejection time. | Tissue Doppler of the mitral annulus from the same subject, showing early peak annular velocity (a measure of peak longitudinal shortening rate), a proxy of volume reduction rate. | Colour tissue velocity from the same subject, with transferred valve openings and closures, showing how early the peak annular velocity is in the ejection time. |
The apex being nearly stationary, the global strain rate (of a wall or the whole ventricle), equals the normalised, inverse value of the matching (wall or the whole ventricle) annular velocity: the annular velocity corresponds fairly closely to the wall strain rate (23).
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As we see, apical velocity is close to zero. | When strain rate (SR) is taken from tissue velocities, the definition is SR= (v(x) - v(x+Δx)) ⁄ Δx where v(x) and v(x+Δx) are velocities in two different points, and Δx is the distance between the two points. If the two points are at the apex and the mitral ring, the apical velocity v(x) ≈ 0, apex being stationary, and v(x+Δx) is annular velocity. Δx then equals wall length (WL), and peak systolic SR = (0 - S') ⁄ WL= (-S') ⁄ WL. |
If the two points are at the apex and the mitral ring, the apical velocity 
, apex being stationary, and 
is annular velocity. 
then equals wall length (WL),
thus 
and peak 
. It's also evident that the basal velocity curve and the strain rate curve approaches each other's shape when strain rate is sampled from most of the wall length. Thus, a method for peak systolic strain rate is peak annular velocity normalised for wall length.
This means that global peak systolic strain rate and global annular peak systolic velocity are physiologically equivalent.
In the HUNT study, the lateral wall had higher S' than the septum, but the global average of peak annular velocity was the same for two walls (septum and lateral wall) as for four walls (septum, anterior, lateral and inferior wall). The differences between walls in strain rate, on the other hand, was minimal.
As shown here, peak rate of volume decrease is reflected in both peak systolic velocity and strain rate in this case, as septal and lateral peak velocities and strain rates are simultaneous.
Comparison between colour Doppler strain rate from the septum (left), and spectral tissue Doppler from the mitral ring of the same subject, showing the same curve shape relations as the colour Doppler derived differences as shown above.
Peak systolic measures are still the closest imaging measures of peak ventricular performance, and are basically
Peak flow is also a measure of peak rate of volume decrease, but with Doppler, we measure flow velocity, while flow = velocity x Area. Thus the flow velocity - flow relation is inveresely related to area. Timing of peak flow, however, should be simultaneous to peak S' and SR.
They are positively related to:
And negatively related to
Thus, peak tissue velocity and peak strain rate are all measures close to peak ventricular performance, but load dependent, and occurring after peak rate of force idevelopment, but before peak myocardial tension.
However, they are both early ejection measures of myocardial performance closer to the peak dP/dt and peak myocardial tension, and are measures of the acceleration of blood out of the ventricles.
Peak systolic annular velocity was early related to ventricular function (123 - 126). Peak strain rate was likewise related to inotropy (127 - 129), and correlated better than strain. Weidemann (127 - 128) did not find that SR related to HR, but this might be that HR was increased by pacing, and thus to reduced ventricular filling, and thus concomitant reduction in preload.
In an experimental study with positive and negative inotropy, Greenberg found a correlation between LV elastance and both S' and peak SR (129).
In a human study with Dobutamin, Thorstensen (130) found that all early systolic indices were more sensitive to changes in inotropy, with a 50 - 60% increase in all three measures with Dobutamine (compared to a 13% increase in HR and 14% in SBP), and with an 11 to 15% decrease (compared to a 13% decrease in HR and 22% decrease in SBP) with Metoprolol. None of the responses of the peak systolic indices were significantly different from the others.
Thus, the peak is not only a function of acceleration, but also related to the early cut off of peak flow rate, due to the increase in aortic pressure.
Greenberg found a stronger correlation of LV elastance with peak SR, than with S' (129). This, however, may be due to the open chest model, which increases translational artefacts, which would affect velocities more that strain rate.
In the HUNT study, using the segmental length method by tissue Doppler tracking:
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Serch areas for kernel tracking from frame to frame, oraqnge, lo0ngitudinal search areas by tissue Doppler, white areas transverse serch areas for speckle tracking. | Real time tracking of kernels at the segment borders. | Strain rate curves. Green: average of three segments of the wall, blue, curve for each segment |
In addition, we examined the peak systolic annular tissue Doppler values by spectral Doppler (S'), and the S' normalised for wall lengths (19) as an equivalent of strain rate as shown above.
Both strain rate and myocardial velocities were normally distributed (19), and so war S'RV.
The normal values for strain rate and and annular velocities by age and sex were (16, 17, 23):
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.
S'LV, mean of 2 walls (cm/s) | S'LV, mean of 4 walls (cm/s) | S'nLV, mean of 2 walls (s-1) | S'nLV, mean of 4 walls (s-1) | SRLV mean of 6 walls (S-1) | Peak LVOT (m/s) | S'RV (cm/s) | ||
(pw TDI) | (pw TDI) | cTDI | (pw TDI) | (pw TDI) | Segmental TDI | Single site | ||
Females | ||||||||
< 40 years | 8.8 (1.1) | 8.9 (1.1) | 7.2 ( 1.0) | 0.94 (0.12) | 0.94 (0.12) | 1.09 (0.12) | 1.01 (0.17) | 13.0 (1.8) |
40 - 60 years | 8.1 (1.2) | 8.1 (1.2) | 6.5 (1.0) | 0.88 (0.13) | 0.88 (0.14) | 1.06 (0.12) | 1.02 (0.16) | 12.4 (1.9) |
> 60 years | 7.3 (1.2) | 7.2 (1.2) | 5.7 (1.1) | 0.81 (0.13) | 0.82 (0.12) | 0.98 (0.14) | 1.01 (0.17) | 11.8 (2.0) |
All | 8.2 (1.3) | 8.2 (1.3) | 6.6 (1.1) | 0.89 (0.13) | 0.89 (0.14) | 1.05 (0.13) | 1.01 0.16) | 12.5 (1.9) |
Males | ||||||||
< 40 years | 9.3 (1.4) | 9.4 (1.4) | 7.6 (1.2) | 0.90 (0.14) | 0.90 (0.14) | 1.06 (0.13) | 0.99 (0.17) | 13.2 (2.0) |
40 - 60 years | 8.6 (8.1) | 8.6 (1.3) | 6.9 (1.3) | 0.84 (0.13) | 0.84 (0.15) | 1.01 (0.12) | 0.99 (0.18) | 12.8 (2.2) |
> 60 years | 8.1 (1.3) | 8.0 (1.3) | 6.4 (1.2) | 0.82 (0.14) | 0.83 (0.13) | 0.97 (0.14) | 0.96 (0.18) | 12.5 (2.3) |
All | 8.6 (1.4) | 8.6 (1.4) | 6.9 (1.3) | 0.85 (0.14) | 0.85 (0.14) | 1.01 (0.13) | 0.98 (0.18) | 12.8 (2.2) |
Total | 8.4 (1.3) | 8.4 (1.4) | 6.8 (1.3) | 0.87 (0.14) | 0.87 (0.14) | 1.03 (0.13) | 1.00 (0.18) | 12.6 (2.1) |
Relative SD (%) | 15.5 | 16.7 | 16.1 | 16.1 | 12.6 | NA | NA | |
Annular velocities and strain rates 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. ll differences between sex and age were significant; all P < .001, but segemntal SR of 2 and 4 walls were not significant, due to too many segments dropped out. Overall standard deviations are given as % of mean in the bottom line, to compare the biological variations between normalized and non-normalized measures. MAPSEn and S′n MAPSE and S′ normalized for LV mean diastolic wall length, respectively. MAPSEn2 and S′n2 normalized for both mean LV diastolic wall length and LV diastolic external diameter, respectively. Abbreviations: GLS = global longitudinal strain; GLSR = global longitudinal strain rate; MAPSE = mitral annular plane systolic excursion; S′ = peak mitral annular systolic longitudinal velocity.
Normal range is customary defined as mean ± 2 SD.
Colour TDI gave lower values than pw TDI as is known since spectral Doppler gives peak values by spectral width, while colour Doppler gives mean values by autocorrelation..
All measures of peak volume change correlated with each other: S' with Global SR by normalised S': unsurprisingly R = 0.87, but alsi with GLSR by segmental tissue Doppler: R = 0.43.
LVOTvmax vs spectral S' : R=0.22, (p<0.001), vs segmental SR 0.17 (0.001), and vs normalised tissue velocity 0.22 (0.001).
But as opposed to SR and S', which declined with age, and correlated negatively respective positively with BSA as described above, LVOTvmax did not correlate with either. As SV is a function of both LVOT velocity (integral) and LVOT area, this indicates that LVOT area is a function of body (heart) size, and that it possibly declines with age. This means that LVOT area is a confounder, and that peak flow velocity in LVOT is not a measure of peak shortening.
- S' declines with increasing age,
- is higher in men than women,
The table shows that:
Both velocity and both methods for strain rate shows significantly decreasing absolute values by increasing age.
S', normalised S' and GLSR versus body size (BSA). S' is weakly positively correlated with BSA, while strain rate (both normalised S' and GLSR) are negatively correlated, and with a larger absolute correlation.
the finding that there was no increase in strain rate by increasing body size (BSA), despite increase in S', relates to the fact as elaborated below, that with a larger BSA, the LV is larger, and the SV is larger. However, a larger LV, means that it is both longer and wider (19). As SV is between 60 and 70% dependent on MAPSE (64 - 68), a wider ventricle, will generate a larger SV, even if MAPSE is unchanged. SR, on the other hand only corrects for LV length, and thus induces a systematic error.
Diagram showing that as diameter (and cross sectional area) is larger in a longer ventricle, given the same MAPSE, SV can be higher with the same MAPSE, but GLS will be lower in absolute values
The peak performance of the LV is seen during the first part of the ejection as described above.
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Pressure-volume-flow diagram showing peak flow in early ejection, peak pressure in mid ejection and peak volume decrease at end ejection. | Annulus velocity and displacement durves, showing peak velocity early after QRS, and peak displacement at end ejection. | Strain rate and strain curves, showing peak (absolute) strain early after ejection, (the peak is less sharp here, as it is from a smaller part of the wall) and peak (absolute) strain at end ejection. |
Velocity, strain rate, displacement and strain. S': peak annular systolic velocity, MAPSE: end systolic Mitral Annular Plane Systolic Excursion, Peak SR: peak systolic strain rate, ESS: End systolic strain. All measures are mean of seoptal and lateral for global measures.
Systolic annular velocity, comparison between HUNT3 and 4
HUNT 3 | HUNT 4 | |||||||||||||
Age group | S' Septal | SEM | 95% CI | S' lateral | SEM | 95% CI | Age group | N | S' septal | SEM | 95% CI | S' lateral | SEM | 95% CI |
Women | ||||||||||||||
< 40 | 8.4 (1.0) | 0.071 | 8.26 - 8.54 | 9.3 (1.6) | 0.109 | 9.08 - 9.52 | 20 - 39 | 64 | 8.5 (1.4) | 0.175 | 8.15 - 8.85 | 10.2 (2.2) | 0.275 | 9.65 - 10.75 |
40 - 60 | 7.7 (1.5) | 0.063 | 7.57 - 7.83 | 8.6 (1.6) | 0.085 | 8.43 - 8.77 | 40 - 59 | 357 | 7.9 (1.3) | 0.069 | 7.76 - 8.04 | 9.4 (2.1) | 0.111 | 9.18 - 9.62 |
> 60 | 6.1 (1.2) | 0.112 | 5.88 - 6.32 | 7.6 (1.5) | 0.137 | 7.33 - 7.87 | 60 - 70 | 355 | 7.2 (1.4) | 0.074 | 7.05 - 7.35 | 8.5 (2.1) | 0.111 | 8.28 - 8.72 |
> 80 | 12 | 6.3 (0.8) | ||||||||||||
Men | ||||||||||||||
< 40 | 8.7 (1.2) | 0.109 | 8.48 - 8.92 | 9.9 (1.9) | 0.172 | 9.56 - 10.24 | 20 - 39 | 49 | 9.4 (1.5) | 0.214 | 8.97 - 9.83 | 11.9 (2.1) | 0.30 | 11.3 - 12.5 |
40 - 60 | 8.1 (1.2) | 0.067 | 7.97 - 8.23 | 9.0 (1.9) | 0.104 | 8.79 - 9.21 | 40 - 59 | 284 | 8.3 (1.5) | 0.089 | 8.12 - 8.48 | 10.0 (2.3) | 0.136 | 9.73 - 10.27 |
> 60 | 7,7 (1.1) | 0.093 | 7.51 - 7.89 | 8.4 (1.7) | 0.144 | 8.11 - 8.69 | 60 - 70 | 279 | 7.9 (1.4) | 0.084 | 7.73 - 8.07 | 9.5 ((2.3) | 0.137 | 9.23 - 9.77 |
> 80 | 12 | 7.4 (1.2) | 8.8 (2.5) | |||||||||||
Both studies show decline with age, but the difference in peak values between the two studies is significant.
However, to assess if this difference is sufficient explanation, we looked at S' isolated in the septum, and still found that the values for HUNT3 and HUNT4 were different for all age groups.
There was a slight difference in mean age, but as the age groups are all significantly different, this is not a reason for the differences.
HUNT 3 was acquired in 2006 - 2008, HUNT 4 in 2017 - 2019. Thus the echo populations are two different cohorts. The two cohorts were similar in sex distribution, BMI and BP.
Both normal studies excluded patients with heart disease, diabetes and hypertension.
HUNT3 was taken on GE Vivid 7, and analysed in EchoPAC version BT06, (except strain and strain rate, which was analysed in the proprietary segmental strain analysis software.
HUNT4 was acquired on GE Vivid E95 and analysed on EchoPAC version 203.
The peak values of spectral Doppler are gain dependent, but this was known at the time of HUNT3, so both studies were acquired with the same convention for spectral Dopppler:
- Peak top of the spectrum
- at lowest possible gain not creating drop outs.
But as there were 10 years technical development, image quality may have improved, enabling a lower gainj setting, and thus lower values in HUNT4
there may be other differences as well:
Package size
Analysis window
Use of sliding windows technique
..............
The main point is that older normal values are not applicable to newer hard- and software.
Systolic velocities vary according to site, as shown in the diagram below.
Site variation of systolic velocites (mean values from HUNT3
This is discussed in more details in the chapter of regional MAPSE, but is due to (156):
The
Age | Septal | Lateral | Anterior | Inferior | LV, mean of sep&lat | LV, mean of 4 walls | S'RV |
Females | |||||||
< 40 years | 8.4 (1.0) | 9.3 (1.6) | 9.2 (1.2) | 8.8 (1.2) | 8.8 (1.1) | 8.9 (1.1) | 13.0 (1.8) |
40 - 60 years | 7.7 (1.5) | 8.6 (1.6) | 8.2 (1.8) | 8.3 (1.3) | 8.1 (1.2) | 8.1 (1.2) | 12.4 (1.9) |
> 60 years | 6.1 (1.2) | 7.6 (1.5) | 7.0 (1.7) | 7.4 (1.4) | 7.3 (1.2) | 7.2 (1.2) | 11.8 (2.0) |
All | 7,78 (1.2) | 8.6 (1.6) | 8.2 (1.8) | 8.3 (1.3) | 8.2 (1.3) | 8.2 (1.3) | 12.5 (1.9) |
Males | |||||||
< 40 years | 8.7 (1.2) | 9.9 (1.9) | 9.5 (2.0) | 9.4 (1.4) | 9.3 (1.4) | 9.4 (1.4) | 13.2 (2.0) |
40 - 60 years | 8.1 (1.2) | 9.0 (1.9) | 8.3 (1.9) | 8.9 (1.5) | 8.6 (8.1) | 8.6 (1.3) | 12.8 (2.2) |
> 60 years | 7.7 (1.1) | 8.4 (1.7) | 7.7 (1.8) | 8.2 (1.3) | 8.1 (1.3) | 8.0 (1.3) | 12.5 (2.3) |
All | 8.1 (1.2) | 9.0 (1.9) | 8.4 (2.0) | 8.8 (1.5) | 8.6 (1.4) | 8.6 (1.4) | 12.8 (2.2) |
Total | 8.0 (1.2) | 8.8 (1.8) | 8.3 (1.9) | 8.6 (1.4) | 8.4 (1.3) | 8.4 (1.4) | 12.6 (2.1) |
Annular velocities by wall, sex and age. Values are mean (SD). Pulsed Tissue Doppler recorded at the top of the spectrum with minimum gain, differences between sex and age were significant; all P < .001.
The mean of 2 walls (8.37 cm/s) was slightly different from the mean of 4 walls (8.40 cm/s) (p < 0.001 in pairwise T-test), statistically significant, but totally but uninteresting as the limit for the accuracy of spectral Doppler is only 0.1 cm/s.
Age | Septal | Lateral | Estimated mean of sep-lat | Anterior | Inferior | Estimated mean of 4 | Anteroseptal | Inferiolateral | mean of 6 | S' RV |
Women | ||||||||||
20 - 39 | 7.0 (1.0) | 7.8 (1.9) | 7.4 | 8.3 (2.1) | 7.4 (1.1) | 7.6 | 6.2 (1.3) | 7.2 (1.9) | 7.3 (1.2) | 11.2 (1.6) |
40 - 59 | 6.5 (1.0) | 7.4 (1.6) | 7.0 | 7.1 (1.9) | 7.0 (1.4) | 7.0 | 5.7 (1.2) | 7.2 (1.7) | 6.8 (1.0) | 10.8 (1.6) |
60 - 79 | 5.8 (1.1) | 6.8 (1.7) | 5.9 | 6.2 (1.7) | 6.2 (1.7) | 6.3 | 5.0 (1.2) | 6.8 (1.6) | 6.2 (1.1) | 10.4 (2.1) |
> 79 | 5.3 (0.8) | 6.6 (1.1) | 6.0 | 5.7 (1.4) | 5.4 (0.6) | 5.8 | 4.4 (1.1) | 5.9 (1.5) | 5.5 (0.7) | 10.3 (1.9) |
All | 6.2 | 7.2 | 6.5 | 6.8 | 6.6 | 6.7 | 5.4 | 7.0 | 6.6 | 10.6 |
Men | ||||||||||
20 - 39 | 7.8 (1.2) | 9.3 (2.0) | 8.6 | 9.2 (2.0) | 8.7 (1.3) | 8.8 | 7.1 (1.3) | 9.0 (1.9) | 8.6 (1.3) | 11.2 (1.6) |
40 - 59 | 7.0 (1.1) | 8.0 (2.2) | 7.5 | 7.8 (2.3) | 7.8 (1.3) | 7.2 | 6.0 (1.4) | 7.7 (2.5) | 7.5 (1.2) | 11.1 (2.0) |
60 - 79 | 6.6 (1.1) | 7.7 (2.2) | 7.2 | 7.2 (2.1) | 7.2 (1.3) | 7.2 | 5.5 (1.5) | 7.4 (2.1) | 7.0 (1.3) | 11.1 (1.9) |
> 79 | 6.5 (1.5) | 7.3 (2.6) | 6.9 | 5.8 (1.6) | 6.5 (1.2) | 6.5 | 5.2 (1.0) | 7.3 (3.1) | 6.4 (1.5) | 12.0 (1.9) |
All | 6.5 | 8.0 | 7.4 | 7.6 | 7.6 | 7.3 | 5.8 | 7.7 | 7.3 | 11.1 |
Total | 6.3 | 7.5 | 6.9 | 7.1 | 7.1 | 7.0 | 5.6 | 7.3 | 6.9 | 10.8 |
Values were corrected for the numbers in each age class before averaging for the cohort. The age group > 79 had only 12 persons of each sex, so cinsidering them as representative is meaningless.
CTDI inh HUNT 4 were significantly lower than pwTDI. This is due to the point that cTDI, corresponds to the middle of the spectrum in spectral TDI (modal velocity), while pwTDI measures at the top of the spectrum as discussed in the basic ultrasound section.
The HUNT3 measures by pwTDI S' only in four points, But we found no difference between global S' as mean of septal-lateral and mean of septal-lateral-anterior-inferior (23). HUNT 4, by cTDI, measures in all 6 walls, and the paper only gives the mean of all 6. As seen below, there is a difference, as the anteroseptal values are lower, but not enough to influence the global mean significantly, in line with what we found for MAPSE in HUNT 3 (156).
Despite the differences in method of TDI, the two studies, both studies confirms the general findings:
- S' declines with increasing age,
- is higher in men than women,
- is lowest in the septum and highest in the lv in the lateral wall, and highest over all in the RV lateral wall.
The total performance of the ventricles corresponds to the stroke volume, which is the measure of the total systolic volume reduction, and thus the total ventricular performance, and cumulated myocardial shortening, although a closer approximation would include the pressure (pulse pressure, PP) this work is performed in, i.e stroke work, which is W = SV × PP.
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 diameter shortening is defined as
FS = (LVIDd - LVIDs) / LVIDd
Diameter is conventionally measured to the endocardium, so the fractional shortening is more precisely the endocardional fractional shortening. Thus, it is in fact an one-dimensional version of EF, under the assumption of a symmetrical ventricle without regional dysfunction. It's less accurate than the EF when there is regional dysfunction, as the measured fractional shortening will be generalised to the whole ventricle.
Endocardial fractional shortening, is obviously related to wall thickening, as the wall shortens longitudinally, it will thicken transversely, given the low compressibility of the myocardial volume.
Deformation in systole. Left: end diastolic image, showing the end diastolic length (Ld = L0). During systole, the ventricle shortens with 
L, which gives (L = L0 - L = Ls). But in order to conserve myocardial volume, the wall thickens at the same time, as shown by the horisontal arrows.
Wall thickening is
WT = (Ws - Wd) / Wd
However, there is also a systolic outer diameter decrease, contributing to the wall thickening as well as the internal diameter decrease.
In HUNT 3 we found the following short axis changes:
Age (years) | Endocardial FS (%) | Outer FS (%) | WT IVS (%) | WT PW (%) | WT mean (%) |
Women | |||||
< 40 | 36.6 (6.1) | 14.1 (3.3) | 45.8 (25.7) | 77.5 (29.4) | 61.7 (20.2) |
40 - 60 | 36.5 (6.9) | 13.2 (4.2) | 44.6 (23.7) | 71.2 (27.6) | 57.9 (19.6) |
> 60 | 36.0 (9.1) | 12.1 (4.2) | 43.7 (22.6) | 65.2 (30.4) | 54.5 (19.8) |
All | 36.4 (7.1) | 13.3 (4.0) | 44.8 (24.1) | 72.2 (28.9) | 58.5 (19.9) |
Men | |||||
< 40 | 35.5 (6.9) | 12.6 (3.7) | 44.5 (19.9) | 68.3 (29.8) | 56.4 (19.1) |
40 - 60 | 35.8 (7.4) | 12.2 (3.8) | 44.1 (22.6) | 65.2(27.0) | 54.6 (19.7) |
> 60 | 36.0 (8.0) | 11.8 (4.4) | 41.3 (21.1) | 62.2 (23.4) | 51.8 (16.4) |
All | 35.8 (7.5) | 12.2 (3.9) | 43.5 (21.1) | 65.2 (26.8) | 54.2 (18.8) |
Total | 36.1 (7.3) | 12.8 (4.0) | 44.2 (22.7) | 68.9 (28.1) | 56.5 (19.6) |
While endocardial FS did not decrease with increasing age, outer FS did, which is related to the decrease in long axis shortening with increasing age.
FS is preserved, however, LVIDd WT and outer FS are all decreasing with increasing age, son there is a decreasing absolute inner diameter shortening of a decreasing diameter. The results are thus consistent. .
In HUNT 4 (249), values are 2D derived and calculated from basal measures: calculating wall thickening and FS from systolic and diastolic values for LVID, IVS and LVPW. Outer FS by LVEDs and LVEHDd, in turn derived from wall thicknesses and internal diameters. Values were corrected for the numbers in each age class before averaging.
Age (years) | Endocardial FS | Outer FS | WT IVS | WT PW | WT mean |
Women | |||||
20 - 39 | 0.32 | 0.16 | 0.45 | 0.41 | 0.43 |
40 - 59 | 0.31 | 0.15 | 0.39 | 0.41 | 0.39 |
60 - 79 | 0.31 | 0.15 | 0.33 | 0.32 | 0.32 |
> 79 | 0.31 | 0.13 | 0.40 | 0.29 | 0.35 |
All | 0.32 | 0.15 | 0.37 | 0.37 | 0.37 |
Men | |||||
20 - 39 | 0.32 | 0.15 | 0.43 | 0.45 | 0.44 |
40 - 59 | 0.31 | 0.15 | 0.36 | 0.35 | 0.35 |
60 - 79 | 0.32 | 0.16 | 0.29 | 0.31 | 0.30 |
> 79 | 0.28 | 0.13 | 0.24 | 0.39 | 0.27 |
All | 0.32 | 0.15 | 0.34 | 0.34 | 0.36 |
Total | 0.31 | 0.15 | 0.36 | 0.35 | 0.34 |
We see there is little difference between IVS and PW. Also the values for the > 79 years may be due to very low numbers in these groups. Else, there is decrease in wall thickening by age, as in HUNT3.
Interestingly, the M-mode values of HUNT3 showed a substantial higher wall thickening in the PW than in the septum, while the 2D measurements in HUNT 4 did not reproduce this finding. This effect is probably due to the specific vulnerability of M-mode to the effects of the long axis shortening, making the M-mode cress different parts of the LV in systole and diastole. The configuration of the posterior wall in then base may thus induce a statistical bias towards over estimation of wall thickness as shown below.
Immages from different parts of the heart cycle, showing that the line crosses different parts of the LV in end diastole and end systole. As the end systolic frame has moved the base of the heart further towards the apex, and the posterior wall thickens towards the mitral annulus, the motion induces an apperent over estimation of end systolic thickness, which will be reflected in the M-mode measurements:
The apparent higher thickening of the posterior LV, may thus be due to the increased thickness of the posterior wall moving into the M-mode line due to the longitudinal motion of the basal parts of the heart.
Pandian (36) by 2D measurements, found substantial heterogeneity of segmental thickening between segments, from WT 34 in septum to 57 in posterior wall, but not as systematic, and using short axis B-mode, may be vulnerable toi long axis basal motion as described for B-mode-
Comparing with longitudinal deformation of the two walls, we found in HUNT3 that MAPSE was about 14% higher in the posterior wall than the anteroseptum, but the posterior wall was also around 10% longer than the anteroseptum (156). Thus, the relative shortening (longitudinal wall strain) in HUNT 3 was 16.6% in the anteroseptum, vs 16.5% by segmental strain, and 14.7% vs 15.5% (relative difference 5%) by normalised MAPSE.
Thus, as longitudinal shortening and transverse thickening are interrelated as shown above, similar relative longitudinal shortenings between the walls, also indicates similar wall thickenings. Thus, the physiology weighs in favor of HUNT4 in this case, while the longitudinal and transmural deformation data in HUNT3 are somewhat inconsistent.
- Both studies, however, show very little decrease in endocardial FS with age.
- Both studies show decrease in relative wall thickening of both walls by age, which, at least in part is related to the age-dependent increase in wall thickness of both walls, found in both studies.
However, in HUNT4, With unchanged endocardial FS, and decreasing WT, it seems that this should add up to decrease in outer FS, as this is calculted from the basic measured, this must be an effect of increasing wall thickness with age, indicating that absolute wall thickening increased. his is in opposition to HUNT3, and what is true may still not be clear, although the decreasing outer diameter may bin HUNT3 may also be an effect of AV-plane motion as seen above..
Stroke volume can be derived from Doppler flow, as described above:
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Tracing the flow velocity curve by pulsed Doppler in LVOT through one heartbeat, gives the velocity time integral by the area under the curve. The LVOT diameter, can be measured in the B-mode. | The velocity time integral is the distance d, that somethin moving with the velocity of the traced curve moves, the stroke distance. The area A of the LVOT (assuming a circular cross section) is given by the measured LVOTdiameter. Thus, the volume of the cylinder given by d × A, equals the stroke volume. |
SV is also available by volumetry of the LV, both by 2D and 3D methods:
Simplified ellipsoid diagram of the LV, showing the principal relation of systolic and diastolic volumes as well as the relation to the dimension changes.
3D echo with LV volume curve. The volumes are all internal. Mark that the ejection curve is sigmoid, as the blood has inertia, the acceleration of the blood after AVO will take a certain time, and thus the maximum rate of volume decrease (=maximum flow rate) is delayed after AVO.
The delay after AVO follows necessarily after physics, instantaneous acceleration to peak flow is contrafectual, but is evident from LVOT Doppler, and has also been documented by volumetry (265, 266).
With a completely incompressible myocardium, the stroke volume would equal the reduction in outer volume, without taking cavity and wall thicknesses into consideration. As the incompressible myocardial volume remains constant, the outer volume reduction must equal cavity volume reduction as shown below.
The left figure shows the cavity volume reduction, being the function of longitudinal and endocardial transverse diameter shortening. But the right figure shows that the total LV outer volume is the sum of cavity and myocardial volume. Given a minimally incompressible myocardium, the reduction in total volume = reduction in cavity volume, while the myocardial volume is constant.
Total (outer) LV volume LVTV = Cavity volume + myocardial volume (MV).
Diastole: LVEDV = LVEDTV - LVEDMV
Systole: LVESV = LVESTV - LVESMV
SV = LVEDV - LVESV = (LVEDTV - LVEDMV) - (LVESTV - LVESMV) = LVEDTV - LVEDMV - LVESTV + LVESMV
If the myocardium is nearly incompressible is LVEDMV 
LVESMV then SV LVEDTV - LVESTV
Stroke volume systolic outer LV volume decrease
While SV may be the most important result of cardiac pumping, it confers little information about the state of the heart itself. In a dilated heart, the stroke volume relative to the total cavity volume (LVEDV) can be reduced, but the absolute stroke volume may be maintained. Although the dilated LV may have an increased afteroad, according to the law of Laplace:
having a larger radius and thinner wall, the ejected volume in terms of the total volume is lower.
Ejection fraction is SV normalised for LVEDV: EF = SV / LVEDV. It is important to realise that Reduced EF, due to chronic dilation, is not related to the Frank Starling effect due to increased preload. In chronic dilation, the EF is a measure of reduced performance, while both dilated and normal ventricles which dilates as response to increased LVEDP, will show increased SV. However, EF will decrease in response to both acute and chronic increase in afterload.
Stroke volume and EF are thus preload dependent, contractility dependent, afterload dependent.
Thus, the ejection fraction is a measure of the relative performance of the LV, in dilated ventricles:
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, and by normalising the SV for EDV, one comes closer to a relative systolic perfomance. In assessing EF, however, 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.
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Normal ventricle with normal EF | Dilated ventricle with increased LV volume, and sverely reduced EF. |
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, basically because inclusion usually was based on dilated conditions.
When EF is evaluated prognostically in an unselected population, however, when patients without dilation is included, it has been shown to be a poor prognosticator even in heart failure (136). This shows that there are shortcomings in the use of EF as a marker of cardiac performance. 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, and this is the reason for this shortcoming, as will be explained below.
Applying the ellipsoid model to the HUNT3 data, gave the following values:
Age | LVEDV(ml) | SV(ml) | EF(%) | Myocardial volume d (ml) |
Women | ||||
<40 | 111.6(21.6) | 76.3(16.4) | 68(6) | 87.0 (19) |
40-60 | 106.9(21.7) | 72.7(17.0) | 68(6) | 92.8 (19.6) |
>60 | 97.9(19.7) | 65.4(16.9) | 66(9) | 95.6 (18.9) |
Total | 106.8(21.8) | 72.6(17.3) | 68(6) | 91.4 (19.6) |
Men | ||||
<40 | 144.8(30.5) | 96.1(22.9) | 66(8) | 125.3 (23.6) |
40-60 | 138.1(31.1) | 92.2(23.8) | 67(8) | 129.7 (25.3) |
>60 | 126.3(33.7) | 84.1(25.7) | 66(8) | 128.2 (26.8) |
Total | 136.6(32.2) | 91.0(24.4) | 67(8) | 128.4 (25.3) |
All | 121.1(31.1) | 81.4(22.9) | 67(8) | 101.4 (27.9) |
LV volumes in HUNT4
Volumes in HUNT 4 are 2D measurement derived, values given are taken from (249).
Comparing with the values from HUNT4 (249), which were measured in 2D, the values according to age and sex can be found in the original publication.:
Age (years) | LVEDV(ml) | SV(ml) | EF(%) |
Women | |||
20 - 39 | 114 (26) | 68 | 60 |
40 - 59 | 102 (19) | 62 | 61 |
60 - 79 | 84 (19) | 51 | 60 |
> 79 | 67 (7) | 41 | 62 |
All | 94 | 57 | 61 |
Men | |||
20 - 39 | 145 (28) | 84 | 58 |
40 - 59 | 136 (29) | 81 | 60 |
60 - 79 | 119 (27) | 71 | 60 |
> 79 | 104 (18) | 62 | 59 |
All | 128 | 76 | 59 |
Total | 109 | 66 | 60 |
SV is calculated from EDV and ESV, corrected for the numbers in each age class before averaging.
Compared to the ellipsoid model from M-mode derived values, both volume, SV and EF is lower, but the relation to age, with declining values in increasing age are the same. This of course follows from the simultaneous age dependent decrease in both LVIDd and LVILd by age in both studies. Myocardial volume is not calculated in HUNT4.
Finally, both studies show concomitant decrease in LVEDV and SV by age, which will maintain EF without any kind of cross sectional increase in contraction.
The NORRE study shows even lower EDV and SV (calculated from end systolic and end diastolic volumes given in the paper), but EF intermediate between HUNT 3 and 4. This study also shows that LVEDV decrease with age, while EF is stable. This, of course means that SV must decrease, (being SV = EDV × EF).
Heart size is obviously related to body size. In the HUNT3 study, LVEDV and SV were both related to body size (R = 0.70 and 0.50, respectively, both p<0.001) (65). EF, being the ratio of the two, is nearly not correlated with BSA at all and the weak correlation is negative (R = - 0.07, p < 0.01). Thus, EF is the only volumetric index that is BSA independent.
Diagram showing a strong correlation of both EDV and SV with body size, while there is only a very weak correlation of EF (being the ratio between the two) and body size.
As the base of the heart moves towards the apex, and the apex is stationary, the LV shortening (in absolute units, e.g. cm), must equal the motion of the LV base, i.e. the Mitral Annular Plane Systolic Motion (MAPSE).
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As the apex is stationary, as shown by the upper line, the total systolic LV shortening is equal to the mitral annulus systolic motion towards the apex. | Mitral annulus motion can be assessed by the longitudinal M-mode through the mitral ring, and the total systolic mitral displacement - Mitral Annular Plane Systolic Excursion - MAPSE, equals LV systolic shortening. |
Global MAPSE is the mean of MAPSE measured in the septal and lateral point, or the mean of septal, anterior, inferior and lateral points or from all 6 walls.
MAPSE varies between sites in the LV (156, 221, 222):
Mean AV-plane systolic displacement at different sites, showing the variability in the Mitral annulus, and TAPSE in the tricuspid annulius which is higher than any site in the mitral annulus.
We compared MAPSE from mean of 2 (septum and lateral wall), 4 (septum, lateral, anterior and inferior walls) and 6 (septum, lateral, anterior and inferior, ateroseptal and inferolateral walls).
However, even if significant in a large study, differences between two, four and six walls were less than 1 mm, which is the limit for measurements in a simgle measure, so the averages were in practice identical. As we found less than 1 mm difference in MAPSE between the averages of 2, 4 and 6 walls in HUNT3, it seems that the number of planes are not of consequence for the mean values.
Age | N | MAPSE 4 walls | SEM | 95% CI |
Women | ||||
< 40 | 208 | 1.73 (0.20) | 0.013 | 1.704 - 1.756 |
40 - 60 | 336 | 1.58 (0.23) | 0.013 | 1.554 - 1.606 |
> 60 | 119 | 1.33 (0.22) | 0.020 | 1.29 - 1.37 |
All | 663 | 1.58 (0.26) | ||
Men | ||||
< 40 | 126 | 1.72 (0.22) | 0.020 | 1.69 - 1.76 |
40 - 60 | 327 | 1.58 (0.22) | 0.012 | 1.556 - 1604 |
> 60 | 150 | 1.45 (0.21) | 0.017 | 1.416 - 1.484 |
All | 603 | 1.58 (0.24) | ||
Total | 1266 | 1.58 (0.25) | ||
Relative SD (%) | 16 | |||
MAPSE is given with the mean for each age and sex group, the SEM is calculated, and the 95% CI for each group will then be mean ± 2 SEM.
Global strain is LV shortening normalised for LV length or wall length as explained here. The table shows global strain as MAPSE normalised for mean wall length (MAPSEn) and GLS by segmental strain, which is another linear method. Normalised MAPSE was done by normalising for wall length (or the straight line approximation of that). This means a skewed line slightly longer than the mid cavity straight line as discussed here. Using an ellipsoid model of the LV, calculated mean LV mid cavity length was 92.4 mm external, and 88.8 mm internal length. Mean MAPSE was 15.8 mm, which would result in a relative LV shortening (as opposed to wall shortening) of 17.1% using the external diameter, and 17.9% using internal diameter.
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Normal distribution of MAPSE; skewness 0.003 | minimal positive correlation with BSA (R= 0.12, p<0.001) | negative correlation with age (R = - 0.50, p < 0.001) |
SV and EF are calculated from the ellipsoid LV model. SV, MAPSE, normalised MAPSE and GLS as well as TAPSE all declined by age. EF did not, as explained below, because of the concomitant decrease in LVEDV and SV, VTI did not for the same reason as peak LVOT flow velocity: As SV is a function of LVOT velocityintegral and LVOT area, this indicates that LVOT area possibly declines with age.
In the HUNT3 study, MAPSE was normally distributed, weakly correlated with body size, but in linear regression, sex differences was not significant. MAPSE was somewhat more strongly negatively correlated with age (18). In HUNT 4, Ventricular lengths and GLS were likewise normally distributed. Relation between relative shortening and GLS by speckle tracking, is discussed elsewhere.
Comparing with the values from HUNT4 (245, 249), which were measured in 2D the values according to age and sex can be found in the original publications. The LV lsystolic and diastolic lengths are presumably from volumetry, and thus are the internal lengths. SV is calculated from the difference in 2D systolic and diastolic volumes, EF both as given in (249) and calculated from LVEDV and SV. Measurements corrected for the numbers in each age class before averaging. Systolic and relative shortening are calculated from the basic length measurements, and likewise corrected for numbers. GLS are values from speckle tracking measurements in (245).
Volumetry by Simpson's method. It also gives internal LV lengths, which vcan be used for LV shortening calculations.
Age (years) | N | LVILd-4ch (cm) | SEM | LVILs4ch (cm) | SEM | Syst shortening-4ch | estimated SEM | 95% CI |
Women | ||||||||
20 - 39 | 64 | 8.5 (0.6) | 0.075 | 7.2 (0.5) | 0.063 | 1.3 | 0.138 | 1.024 - 1.576 |
40 - 59 | 357 | 8.3 (0.6) | 0.032 | 7.1(0.5) | 0.026 | 1.2 | 0.058 | 1.084 - 1.316 |
60 - 79 | 355 | 7.9 (0.6) | 0.032 | 6.8 (0.5) | 0.027 | 1.1 | 0.059 | 0.982 - 1.218 |
> 79 | 12 | 7.1 (0.5) | 6.4(0.4) | 0.7 | ||||
All | 8.1 | 8.2 | ||||||
Men | ||||||||
20 - 39 | 49 | 9.7 (0.6) | 0.086 | 8.1(0.5) | 0.071 | 1.6 | 0.157 | 1.286 - 1.914 |
40 - 59 | 284 | 9.2 (0.6) | 0.036 | 7.9 (0.5) | 0.030 | 1.3 | 0.066 | 1.168 - 1.432 |
60 - 79 | 279 | 8.9 (0.6) | 0.036 | 7.7 (0.5) | 0.030 | 1.2 | 0.066 | 1-068 - 1.332 |
> 79 | 12 | 8.5 (0.5) | 7.2 (0.5) | 1.3 | ||||
All | 9.1 | 9.2 | ||||||
Total | 8.5 | 8.6 | ||||||
SEM is calculated from the numbers in each age group, and the SD of the diastolic and systolic lengths, given in the paper. SEM of LV shortening (difference between lengths) is then conservatively assumed to be the sum of the SEM for each length, and the 95% CI is mean ± 2SEM. Comparing with the MAPSE values of HUNT3 in the table above, we see that the 95% CI are non-overlapping, except for the youngest male group, which has a rather wide CI due to the small size.
In general, we must conclude that the two studies show different values for almost all age groups (and thus, the difference is not due to the mean age difference of the two cohorts.
It is reasonable to assume that this is due to the difference in the methods used.
M-mode MAPSE and LV shortening by 2D from the same loop. The two methods give different values as discussed above. We also see that in this case, there is no M-mode artefact of the cursor sliding along the wall, ist is simply that while M-mode reflects external wall shortening, the internal shortening is less.
Thus the two methods are in no way interchangeable. Still, HUNT 4 confirms the decreasing absolute and relative LV shortening with increasing age found in HUNT3, as well as equal absolute, but lower relative shortening in men vs women..
LV shortening was calculated from diastolic and systolic LV lengths given in the paper and supplementary, which are exports from the endocardial traces from volumetry. As we see, both in HUNT 3 and 4, there is little difference in measuring LV shortening in 4-chamber alone or in mean of two planes (4 walls), but the findings of the two studies differ considerably.
The absolute and relative shortening by 2D, are thus less than the M-mode derived values of HUNT 3. This bias points to one or more systematic deviations in the methods.
Using caliper measurements of six points of the Mitral annulus on MR, Ochs et al (222) found AV-plane motion of 15 mm in 209 healthy subjects, in line with HUNT3. Sepulveda-Martinez et al, (251) in a direct comparison between Echocardiography M-mode and caliper based MR measurement in 111healthy subjects found a mean MAPSE of 17 mm, (in a younger cohort), and a bias of only 1 mm between the two methods. Thus, it seems that M-mode of the mitral plane is the most reliable, and that the HUNT3 describes the absolute LV shortening closer.
MR studies in gereal measures outer contour shortening (66).
The main reason is that the internal shortening is less that the external shortening, due to the dome shape of the LV. The mitral ring is a stiff structure, there is no Torsion during systole, and thus the ring motion follows the outer contour (wall) shorteni ng), despite being measured in the cavity.
Internal (blue arrows) and external (orange arrows) LV shortening, showing that the external shortening is larger than the internal, due to the dome shape of the LV basis.. The M mode conforms with the external wall shortening, (even when measured internally, as the mitral ring is a rigid structure, and there is no torsion of the ring during systole.
Age (years) | SV (ml) | EF (%) | LVOT VTI (cm) | MAPSE (4 walls cm) | MAPSEn (4 walls %) | Segmental GLS (%) | TAPSE (cm) |
Women | |||||||
< 40 | 76.3 (16.7) | 68 (6) | 20.8 (3.5) | 1.73 (0.20) | 18.1 (2.0) | 17.9 (2.1) | 2.9 (0.5) |
40 - 60 | 72.7 (17.0) | 68 (7) | 21.6 (3.4) | 1.58 (0.23) | 17.0 (2.2) | 17.6 (2.1) | 2.7 (0.5) |
> 60 | 65.4 (16.9) | 66 (9) | 21.7 (3.7) | 1.33 (0.22) | 14.8 (2.1) | 15.9 (2.4) | 2.5 (0.56) |
All | 72.6 (17.3) | 68 (8) | 21.4 (3.5) | 1.58 (0.26) | 17.0 (2.4) | 17.4 (2.3) | 2.8 (0.5) |
Men | |||||||
< 40 | 96.1 (22.9) | 66 (8) | 20.0 (3.3) | 1.72 (0.22) | 16.5 (2.0) | 16.8 (2.0) | 3.0 (0.6) |
40 - 60 | 92.2 (23.8) | 67 (8) | 20.4 (3.6) | 1.58 (0.22) | 15.4 (1.9) | 15.8 (2.0) | 2.9 (0.6) |
> 60 | 84.1 (25.7) | 66 (9) | 20.3 (3.7) | 1.45 (0.21) | 14.9 (1.9) | 15.4 (2.4) | 2.8 (0.6) |
All | 91.0 (24.4) | 67 (8) | 20.3 (3.6) | 1.58 (0.24) | 15.5 (2.0) | 15.9 (2.3) | 2.9 (0.6) |
Total | 81.4 (22.9) | 67 (8) | 20.8 (3.6) | 1.58 (0.25) | 16.3 (2.4) | 16.7 (2.4) | 2.8 (0.5) |
Relative SD (%) | NA | NA | NA | 16 | 14 | 24 | NA |
Global strain is LV shortening normalised for LV length or wall length as explained here. The table shows global strain as MAPSE normalised for mean wall length (MAPSEn) and GLS by segmental strain, which is another linear method. Normalised MAPSE was done by normalising for wall length (or the straight line approximation of that). This means a skewed line slightly longer than the mid cavity straight line as discussed here. Using an ellipsoid model of the LV, calculated mean LV mid cavity length was 92.4 mm external, and 88.8 mm internal length. Mean MAPSE was 15.8 mm, which would result in a relative LV shortening (as opposed to wall shortening) of 17.1% using the external diameter, and 17.9% using internal diameter.
Age (Years) | SV(ml) | EF (measured %) | EF (calc %) | Syst. shortening (cm) 4-ch | Syst. shortening (cm) 2-ch | Mean shortening (cm) | Relative shortening (%) |
Women | |||||||
20 - 39 | 68 | 60 (4) | 60 | 1.3 | 1.3 | 1.3 | 15.2 |
40 - 59 | 62 | 61 (5) | 61 | 1.2 | 1.2 | 1.2 | 14.4 |
60 - 79 | 51 | 60 (5) | 61 | 1.1 | 1.1 | 1.1 | 13.9 |
> 79 | 41 | 62 (3) | 57 | 0.7 | 0.9 | 0.8 | 11.1 |
All | 57 | 60 | 61 | 1.2 | 1.2 | 1.2 | 14.2 |
Men | |||||||
20 - 39 | 84 | 58 (6) | 60 | 1.6 | 1.4 | 1.5 | 15.5 |
40 - 59 | 81 | 60 (5) | 60 | 1.3 | 1.4 | 1.4 | 14.5 |
60 - 79 | 71 | 60 (5) | 60 | 1.2 | 1.1 | 1.2 | 12.9 |
> 79 | 62 | 59 (5) | 60 | 1.3 | 0.9 | 1.1 | 13.0 |
All | 76 | 59 | 60 | 1.3 | 1.3 | 1.3 | 13.8 |
Total | 66 | 60 | 60 | 1.2 | 1.2 | 1.2 | 14.0 |
Mean values for each group are taken from (249) for lengths, and corrected for the numbers in each age class before averaging. Absolute and relative shortening are calculated from the basic measurements, and likewise corrected for numbers.
In HUNT3, the mean left ventricular internal length was calculated from wall lengths to 88 mm. In HUNT4, the internal LV length was 85 mm.
GLS is by speckle tracking, from (245), which is the same population (for comparison with absolute and relative shortening, GLS is given by numerical values):
<40 years | 40 - 49 years | 50 - 59 years | 60 - 69 years | > 69 years | All |
Women | |||||
21.3 (1.8) | 21.0 (2.2) | 20.5 (1.8) | 19.7 (2.0) | 19.0 (2.0) | 20.2 (2.0) |
Men | |||||
19.8 (1.9) | 20.1 (1.8) | 19.5 (1.8) | 18.9 (1.9) | 18.5 (2.1) | 19.3 (2.0) |
Total: 19.8 (2.1) | |||||
TAPSE was likewise normally distributed, and correlated negatively with age, positively with BSA, small gender differences were not significant in linear regression with BSA (103).
Both TAPSE and S'RV correlated modestly with BSA, and there was a sex difference, but this was simply due to BSA difference, as shown by linear regression.
The eggshell model of the heart would predict that the stroke volume would be solely the function of long axis shortening (12 - 14), at least with an incompressible myocardium. As discussed in the basics section, however, there is an outer diameter decrease as well (62, 63, 65), contributing to stroke volume. This can be seen here:
A>s seen here, the main deformation of the ventricles, is in the longtitudinal direction by AV-plane motion. Most transverse chamber reduction,m is due to wall thickening, which again is a function of shortening. But there is also an outer diameter reduction due to true circumferential fibre shortening, as discussed here, and this outer diameter reduction can be seen here.
With a completely incompressible myocardium, the stroke volume would equal the reduction in outer volume, without taking cavity and wall thicknesses into consideration. As the incompressible myocardial volume remains constant, the outer volume reduction must equal cavity volume reduction as shown below.
The left figure shows the cavity volume reduction, being the function of longitudinal and endocardial transverse diameter shortening. But the right figure shows that the total LV outer volume is the sum of cavity and myocardial volume. Given a minimally incompressible myocardium, the reduction in total volume = reduction in cavity volume, while the myocardial volume is constant.
Total (outer) LV volume LVTV = Cavity volume + myocardial volume (MV).
Diastole: LVEDV = LVEDTV - LVEDMV
Systole: LVESV = LVESTV - LVESMV
SV = LVEDV - LVESV = (LVEDTV - LVEDMV) - (LVESTV - LVESMV) = LVEDTV - LVEDMV - LVESTV + LVESMV
If the myocardium is nearly incompressible is LVEDMV 
LVESMV then SV LVEDTV - LVESTV
Stroke volume systolic outer LV volume decrease
Outer volume decrease has two components:
Longitudinal component = MAPSE × Mitral annular outer area = MAPSE × 
outer mitral annular diameter / 2
Transverse component which is SV - longitudinal component.
In the HUNT study we used a symmetric, ellipsoid model of the LV, In the HUNT3 ellipsoid LV model, (65), we measured MAPSE and outer ventricular diastolic and systolic diameter, and assumed the mitral annular diameter to be equal to LV outer systolic diameter as shown by the figure above. Thus we calculatedThe SV from the cavity volumes, MAPSE × mitral annular area, to derive the MAPSE part of the SV, considering the remaining decrease to be due to the ourer LV diameter decrease. We found that MAPSE contributed 74.2% of total SV. Circumferential shortening due to OUTER circ. (diameter) shortening, was 12.8%, and must make up the rest, 25.8% of SV.
Although all primary measurements were normally distributed, the volumes were not, indicating that there was a systematic error in the geometrical model. This is reasonable, as the assumption of the model was a symmetric ellipsoid, which is not the case in real life.
Direct measurement by MR have shown that the AV-plane contribution is closer to 60% for the LV, but ca 80% for the RV (66, 67,68), which probably is closer, although a study of LV filling found that systole contributed 70% to ventricular filling (9), which should be equal to the ejected volume unless there is concomitant atrial expansion also.
Relations of MAPSE to SV
Age | MAPSE (cm) | SV(ml) | MAPSE vol(ml) | EF(%) | MAPSE% of SV | Endocardial FS(%) | Outer FS (%) | Wall thickening (%) |
Women | ||||||||
<40 | 1.73(0.20) | 76.3(16.4) | 56.5(9.9) | 68(6) | 75.4(11.9) | 36.6(6.1) | 14.1(3.3) | 61.7 (20.2) |
40-60 | 1.58(0.23) | 72.7(17.0) | 53.3(11.7) | 68(6) | 74.9(13.5) | 36.5(6.9) | 13.2(4.2) | 57.9 (19.6) |
>60 | 1.33(0.26) | 65.4(16.9) | 45.2(10.1) | 66(9) | 72.0(21.9) | 36.0(9.1) | 12.1(4.2) | 54.5 (19.8) |
Total | 1.58(0.26) | 72.6(17.3) | 52.9(11.5) | 68(6) | 74.6(14.9) | 36.4(7.1) | 13.3(4.0) | 58.5 (19.9) |
Men | ||||||||
<40 | 1.72(0.22) | 96.1(22.9) | 70.1(14.9) | 66(8) | 74.9(14.2) | 35.5(6.9) | 12.6(3.7) | 56.4 (19.1) |
40-60 | 1.58(0.22) | 92.2(23.8) | 65.1(14.2) | 67(8) | 72.8(14.8) | 35.8(7.4) | 12.2(3.8) | 54.6 (19.7) |
>60 | 1.45(0.21) | 84.1(25.7) | 60.3(14.7) | 66(8) | 74.9(19.0) | 36.0(8.0) | 11.8(4.4) | 51.8 (16.4) |
Total | 1.58(0.24) | 91.0(24.4) | 64.9(14.7) | 67(8) | 73.8(15.8) | 35.8(7.5) | 12.2(3.9) | 54.2 (18.8) |
All | 1.58(0.24) | 81.4(22.9) | 61.5(13.0) | 67(8) | 75.2(12.8) | 36.1(7.3) | 12.8(4.0) | 56.5 (19.6) |
We have calculated LV shortening from LV systolic and diastolic lengths, corrected for the numbers in each age class before averaging. With the same assumptioon as in HUNT3, that mitral annular diameter is close to LV external systolic diameter (LVEDs = LVIDs + LVPWs + LVIVSs), and the part of SV being LV shortening × 
× (LVEDS/2)2.
Age (Years) | SV(ml) | EF (measured %) | EF (calc %) | Mean shortening (cm) | LV shortening vol (ml) | LV shortening vol (%) |
Women | ||||||
20 - 39 | 68 | 60 (4) | 60 | 1.3 | 28 | 41 |
40 - 59 | 62 | 61 (5) | 61 | 1.2 | 26 | 43 |
60 - 79 | 51 | 60 (5) | 61 | 1.1 | 23 | 45 |
> 79 | 41 | 62 (3) | 57 | 0.8 | 15 | 37 |
All | 57 | 60 | 61 | 1.2 | 25 | 44 |
Men | ||||||
20 - 39 | 84 | 58 (6) | 60 | 1.5 | 38 | 45 |
40 - 59 | 81 | 60 (5) | 60 | 1.4 | 36 | 44 |
60 - 79 | 71 | 60 (5) | 60 | 1.2 | 29 | 40 |
> 79 | 62 | 59 (5) | 60 | 1.1 | 27 | 44 |
All | 76 | 59 | 60 | 1.3 | 33 | 44 |
Total | 66 | 60 | 60 | 1.2 | 29 | 44 |
Mean values for each group are taken from (249) for lengths, and corrected for the numbers in each age class before averaging. Absolute and relative shortening are calculated from the basic measurements, and likewise corrected for numbers.
The findings in HUNT 4 deviates from HUNT3 both in absolute and relative long axis shortening, but also in the LV shortening contribution - both absolute and relative - to SV (despite lower SV).
This is somewhat surprising, and as the M-mode derived values
M-mode vs MR
Absolute og relative SV i Carlsson
While MAPSE correlates only weakly with BSA (R = 0.12, p<0.001), The MAPSE contribution to SV in % does not correlate with BSA at all. Thus, the MAPSE % of SV, being constant across the BSA range, MAPSE contribution in ml correlates strongly with BSA (R = 0.55, mp< 0.001) (65).
Diagrams showing the weak correlation of MAPSE with BSA, but also that the MAPSE percentage contribution to SV is BSA independent, which means, as SV is strongly BSA dependent, so is The MAPSE contribution in absolute volume.
The finding of unchanged EF and FS with increasing age, has been found numerous times (31 - 37), also in the HUNT 3 (65). It has been assumed that this was due to a "compensatory increase in short axis function". This argument was repeated by the finding of decreasing MAPSE with increasing age. However, this is not the case.
In the HUNT study half ellipsoid model of the LV based on linear measures (65), the MAPSE contribution to the SV would be external mitral annulus x MAPSE, while the transverse diameter shortening would contribute the rest. The aim of this study was primarily to study the age related changes in strains in relation to EF and stroke volume, as any systematic error might be assumed to be more or less constant across the study group. We found that EDV and SV declined with age. EF was unchanged by age, the mechanism for this was the concomitant decrease in EDV and SV, preserving the ratio.
MAPSE declined with age as shown previously (7, 18, 65), but not the MAPSE percentage of the SV, again due to the concomitant decrease of both measures.
Thus the preserved EF despite decrease in long axis function is NOT due to any compensatory increase in transverse function, all thre strains decline with age. The preserved EF is due to the simultaneous reduction in EDV, SV and MAPSE.
In the HUNT ellipsoid model, we looked at the changes in dimensions and volumes, SV, EF and FS in relation to ageing (65): The systematic error can be assumed to be more or less the same across age groups, and thus the findings in relation to ageing mif´ght be assumed to be valid.
In the HUNT study, we found in addition near unchanged LVIDD, unchanged FS and increased wall thickness, but with decreased wall thickening (relative).
Thus, there is no "compensatory short axis increase as mechanism for maintained EF.
Methodologically, this shows that EF and FS are not sensitive for age related changes.
The long axis shortening and lengthening are thus related to annular displacement, and the displacement curve is closesly related to the absolute volume curve. The annular velocity is the first derivative of the displacement curve, just as the flow is the first derivative of volume.
Relation of pressures, flow velocity, volumes, annular velocity and displacement. Length and velocity are both related to volume changes, and are the resultant of tension vs. load. Both E and e' are related to tension vs load, and e' is not the cause of early pressure changes.
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Normal ventricle with normal EF | Dilated ventricle with increased LV volume, and sverely reduced EF. |
In dilated ventricles, the increasing LVEDV will cause the EF to decrease, despite normal SV. Thus, the EF is a more sensitive measure than SV. Likewise, the FS will be rediced as LVIDD increases in dilation.
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.
MAPSE will also be reduced in dilated heart failure:
When the left ventricle dilates, both the volume and diameter increases, and the stroke volume can be maintained by a smaller fraction (Ejection fraction) of the total (end diastolic) volume. As the stroke volume is proportional with the MAPSE x outer area, as discussed above, the SV also can be maintained with a lower MAPSE and a larger diameter, and there will be a near linear relation between reduction in EF and MAPSE.
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 (95, 133 - 136). The relation between MAPSE and EF has shown a correlation of 80 - 90%. However, the relation only holds in the spectrum from normal to dilated ventricles, when both are included. In a population of only normal ventricles, this correlation is not present, here MAPSE correlates with SV (137). Finally, there is no correlation of MAPSE with EF in hypertrophic ventricles (138).
While EF is a prognosticator in heart failure with dilatation, the same is not the case in a general heart failure population (131). This is due to the presence of heart failure with preserved ejection fraction HFpEF.
Patient with amyoloidosis and heart failure:
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PSLAX | PSSAX | A4CH |
Evident is the thick wall and the small cavity. Systolic function is not normal, as the small cavity is unable to generate a normal SV, but as both LVEDV and SV are reduced, their ratio (EF) may be normal. MAPSE, however, will bve reduced also here.
MAPSE on the other hand, will be reduced, in line with the reduced SV, despite normal EF (139).
In concentric hypertrophy the LVEDV is reduced, and unable to generate the same SV. But as both LVEDV and SV are reduced, the EF may be preserved (HFpEF). The outer diameter may be increased, unchanged or reduced, depending on the pattern of wall thickness increase. Still the SV is proportional to the MAPSE x area, as discussed above, there will be a linear relation between the decrease in MAPSE and SV.
Concentric hypertrophy reduces the cavity volume, as well as the stroke volume, and the concomitant reduction of both can maintain the ratio between them. 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. EF is a geometrical concept, and only works in some geometries. MAPSE will decrease in relation to the reduction in SV.
The lack of understanding of the geometry of EF, has led to two misconceptions:
It is very common to hear that as longitudinal function decrease with age or hypertrophy or whatever, the short axis function shows compensatory increase. Especially if EF is preserved. This mis conception artises from the confusion between cavity and wall measures. In the HUNT study, LV internal diameter, and fractional diameter shortening did not change significantly) with increasing age, while wall thickness increased, and wall thickening (transmural strain) decreased (19, 7).
Age (years) | LVIDd (mm) | FS (%) | Wall thickness (mm) | Wall thickening (transmural strain %) |
< 40 | 51.3 | 36 | 8.1 | 59.1 |
40 - 60 | 50.9 | 36 | 8.9 | 56.3 |
> 60 | 50.0 | 36 | 9.5 | 53.0 |
The finding of unchanged LVIDD and FS has been confirmed before (32 - 37). Thus, there is no sign of increases cross sectional function in increasing age, even with preserved EF (65).
It has also been maintained as a mechanism for preserved EF in HFpEF, as longitudinal function is shown to be reduced. However, this is not found in direct studies of hypertrophy with HFpEF (80- 82).
The reason for the confusion is the different relations of cavity and wall measures to hypertrophy:
Schematic diagram of M-mode, illustrating the relation between cavity measures and wall measures of function. For simplicity, the ventricle is assumed to be symmetric. Left, a normal ventricle with wall thickness of 1 cm, systolic wall thickening of 50% and a systolic 11% outer diameter reduction. This gives a fractional shortening of 36%. Right a concentric hypertrophic ventricle with reduced myocardial function. The wall thickness is increased to 1.5 cm, outer diameter is the same, outer diameter shortening is reduced to 8.6%, wall thickening, while nearly the same in absolute measurement, is now only 30% relative. Fractional cavity diameter shortening, however, is slightly increased to 37.5%, but this is a function of the diastolic diameter reduction, nothing else. The true short axis function is the wall thickening, which is reduced.
So in this case, the short axis function is reduced, not increased, and the apparent increase in FS is in fact a decrease in both EDD, absolute shortening, and wall thicke4ning, preserving or increasing their ratio, analoguous with the volumes in EF.
The above example is an illustration, showing that in some instances, especially hypertrophy, the cavity measures do not measure the myocardial function, and the apparent "increased short axis function" is illusory, based on a misconception. An example illustrates this.
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Patient with heart failure. EDV 100 ml, EF 55%, thus SV 55 ml. | Wall thickness 17 mm, EDD 40 mm. | FS 35%, however, wall thickening 28% |
Thus, in HFpEF, both cavity and LVIDD are reduced, while EF and FS are preserved.
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MAPSE = 5 mm | S' = 3 mm |
The annular displacement has been shown to be more sensitive than EF in predicting events in heart failure (140, 141) and hypertension (142). Also, the MAPSE correlates better with BNP in heart failure, that the fractional shortening (143).
Looking at the systolic AV-plane displacement, the displacement varies intra individually with the site (156), in the LV highest laterally, both by M-mode (221) and by MR (222), but higher in the RV free wall than the LV lateral wall, (156) also found by others (223, 224). This is true also for S', as well as for wall lengths (16), so velocity and diasplacement varies concordantly (156). Looking at wall lengths, we have previously fond that in the LV the variation followed the same pattern (19).
Example of intra individual variability of LV wall lengths (measured as the straight line from the mitral annulus to the apex), annular plane displacement by M-mode, and S'. All are examplified by the septal and lateral measures, showing that both displacement and S' are highest in the RV free wall, lowest in the septum, higher again in the LV lateral wall than the septum. In the LV, the lateral wall is longer than the septum.
Mean AV-plane systolic displacement and velocities at different sites, showing that the highest displacement and velocity in the LV is in the lateral annulus, lowest in the septum, and the S' and TAPSE in the tricuspid annulius is higher than any site in the mitral annulus.
Looking at Mitral and tricuspid annular velocities (16), MAPSE/TAPSE (156) and LV strain rate and strain (17) per wall, the pattern is fairly consistent.
As all walls in the LV works at the same pressure (load), differences in wall shortening must reflect differences in contractility, i.e shortening vs load must be the same.
This, however, is not the same for the RV, where load is considerably lower, and the wall is longer.
Septal | Anteroseptal | Anterior | Lateral | Inferolateral (posterior) | Inferior | Relative, intraindividual variance (var / mean) | |
WL (cm) | 9.2 (1.7) | 9.2 (1.9) | 9.5 (1.8) | 9.6 (1.8) | 10.1 (2.1) | 9.5 (1.8) | 0.4 |
Displacement (MAPSE) (cm) | 1.5 (0.3) | 1.4 (0.4) | 1.5 (0.3) | 1.6 (0.3) | 1.6 (0.4) | 1.7 (0.3) | 0.03 |
MAPSE / WL (%) | 16.2 (2.8) | 14.7 (3.6) | 15.9 (2.8) | 16.3 (2.7) | 15.5 (3.6) | 17.1 (3.0) | 0.003* |
S' (cm/s) | 8,0 (1.2) | 8.3 (1.9) | 8.8 (1.8) | 8.6 (1.4) | 0.1 | ||
S' / WL (s-1) | 0.85 (0.13) | 0.85 (0.2) | 0.90 (0.19) | 0.88 (0.14) | 0.87 (0.14) | 0.01* |
Population means, and population standard deviations in parentheses. *p < 0.001 for difference of relative variance of normalised versus non normalised measures.
Normalising MAPSE and S' for wall length, reduced the relative intra individual variance by a factor of 90%, showing clearly that the systolic displacement and systolic displacement velocity were clearly related to wall lengths. As all walls of the LV works with similar afterload, this shows that contractility (i.e. shortening per length, is similar in all walls), related to the number of sarcomeres in series.
This is evident as also S' is related to wall length, i.e. the regional velocity / acceleration is a function of force.
Normalised MAPSE is a measure of strain (18), normalised S' a measure of strain rate (23).
Segmental strain and strain rate can also be seen to have far less variability than displacement and S', but we didn't react to this in the original studies (16, 17).
Thus, the LV motion is greatest in the lateral parts where the wall is longer, and shortening relative to wall length (wall strain) is more constant, consistent with a relatively uniform longitudinal sarcomere shortening in the LV.
Normal annular peak systolic velocities, strain rate and strain per wall in the HUNT study by tissue Doppler.
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) |
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.
Colour Doppler traces of velocity, displacement, strain and strain rate from the septal (yellow) and lateral (cyan) aspect of the four chamber view. motion traces are from the base,, deformation traces are from the whole wall as shown by the regions of interest (ROI). Systolic motion is positive, towards the probe. Systolic strain rate and strain is negative, as they represent shortening, and this is also evident from the definition of the velocity gradient / Strain rate. From this diagram, it is also evident that the velocities and displacements are higher in the lateral wall than the septum, while strain rate and strain are nearly equal. This is due to the fact that wall strain rate and strain basically are annular velocity and displacement normalised for wall length, and the lateral wall is longer than the septum.
The LV is often shown as a symmetric half-ellipsoid, but in reality, it is a skewed ellipsoid with the apex aligned with the centre of the AV - plane and the septum, which thusis shorter and straighter than the lateral wall.
The relation of cardiac anatomy to regional AV-plane motion. The figure shows the longer walls in the lateral aspects, and shorter walls near the septum, and this relates to the position of the apex over the central part of the AV-plane. The central part of the AV-plane also is the site of the large arteries, which may anchor the AV-plane in the middle. The mean AV plane motion in the different sites are indicated, showing correspondence with WL in the LV.
The large arteries are anchored in the centre of the AV‐plane and must stretch during the systolic AV‐plane motion, so the differences in AV‐plane motion are related both to wall length and to the location of the large arteries, and hence, to the overall anatomy.
Shortening of the RV is greater than the LV, but RV wall is thinner and longer and works against a lower afterload, and hence tension, so the shortening relation to wall length it is not comparable with the LV. As RV motion is only measured in one wall, there are no other walls with similar load for comparison, and a direct comparison of shortening vs length between RV and LV is not relevant, but both interacts with the AV-plane.
Thus, the AV‐plane not only move towards the apex, but also bends into a U‐shape during systole, so not only the ventricular length, but also the width of the AV‐plane and the maximal width of the ventricles are reduced as well as tilts towards the left as shown below.
Example from a single subject. As seen both AV-plane systolic motion and AV-plane peak systolic velocity are highest in the rght lateral part, and lowest in the central part. This means that the differential motion will correspond to a systolic bending of the AV-plane, and a tilting towards the left.
What are the functional significance of the systolic AV-plane bending?
The systolic bending is evident also from the systolic velocities, but from this table it is also evident that most of the unbending happens in early diastole (highest e' in the RV, lowest in the septum), while the av-plane remains straighter, but with a tilting towards the right during atrial systole (highest in the RV, lowest in the left lateral)
Normal annular peak S', e' and a' per wall in the HUNT study by tissue Doppler.
Right ventricle | Anterior | septal | Inferior | lateral | |
PwTDI S' (cm/s) | 12.6 (2.1) | 8.3 (1.9) | 8.0 (1.2) | 8.6 (1.4) | 8.8 (1.8) |
PwTDI e' (cm/s) | 12.9 (3.2) | 11.6 (3.7) | 9.9 (2.9) | 11.2 (3.5) | 12.5 (3.7) |
PWTDI a' (cm/s) | 14.3 (3.8) | 9.4 (2.3) | 10.2 (2.2) | 11.0 (2.3) | 9.4 (2.4) |
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In a modified eggshell model, there is also a reduction in outer transverse diameter, due to both outer ventricular diameter reduction and to AV-plane bending. The AV-plane apical motion is responsible for most of the ventricular systolic volume compression (red) and thus most of the the ejection volume, but in addition the volume is reduced by outer circumferential fibre contraction, and AV plane compression by bending, causing reduction in transverse diameter adding to the volume reduction (orange) and ejection. The atrial expansion, however, and thus the systolic venous inflow (S-wave), is mostly equal only to that caused by AV-plane motion. | During early filling there is recoil of the AV-plane, but only partly back to the end diastolic position, causing ventricularisation of part of the atrialised volume behind the compressed AV-plane (light red). This effect do not generate flow. The unbending of the AV-plane will cause an additional reduction of the atrial volume (violet), causing pressure rise (the V-wave), and atrioventricular early flow. | Atrial contraction will empty the auricles, giving an additional volume (light blue) that is injected into the ventricles, causing the final ventricular expansion back to the end diastolic position. This injection of blood causes pressure rise in atrium, |
Below is a clinical example corresponding to the diagrams above:
The three main coordinates of the heart: longitudinal, transmural and circumferential. This is the normal strain tensors, in this coordinate system, i.e. the coordinates of the deformation.
Global strain is basically left ventricular shortening normalised for end diastolic length. As discussed here, there is no universal standard, as the numerator depends on the method for tracking, and the denominator on what is chosen for reference length.
Inward tracking, is a property of feature tracking, and will result in apparent longitudinal shortening of the midwall and endocardium, even, hypothetically, without any ventricular shortening at all. This is due to the conical shape of the LV.
However, the common ground is that it is the left ventricular shortening, normalised in some way by LV end diastolic length. As MAPSE related to stroke volume, so did GLS.
It was shown early in experimental studies, that while peak strain rate was more closely related to contractility (ventricular elastance and inotropy) , end systolic strain and MAPSE was more closely related to SV and EF (127 - 130). The close relation of LV volume change and strain is shown below:
The picture shows a detailed LV volume curve from a healthy person by MUGA scintigraphy, the left a normal longitudinal strain curve. The similarity shows:
the volume changes of ejection and filling are closely related to the changes of the LV longitudinal dimension. However, as the left shows volume changes, while the right shows wall length changes, and as wall length changes vary between walls, the correspondence is not perfect.
In the HUNT3 study, using the geometrical model, Global MAPSE correlated woth normalised global MAPSE (R=0.86), GLS by segmental strain (R=0.40), S' (R=0.34), SV (R=0.25) and EF (R=0.16), all p<0.001). The two methods for GL correlated with each other (R=0.52), with S' (R=0.26 and 0.44, respectively), with EF (R=0.22 and 0.24), all p<=.001). However, GLS by either method do not correlate with SV (156). The possible explanation is the previous finding that SV is related to both LV length and diameter, which are interrelated, while global LV strain only normalizes for LV length, thus introducing a systematic error as described above. The SV in this comparison, however, is derived from a geometrical model (65), which in itself may include a systematic error, but the concordant results between the two strain methods, as well as the maintained relation of SV with MAPÅSE and S’, supports this finding.
Both segmental tissue Doppler derived GLS, normalised MAPSE and MAPSE itself, were normally distributed.
Normal distributions of MAPSE, MAPSEn, and GLS, with skewnesses of 0.003, -0.17 and 0.19, respectively.
It does seem intuitive that normalising MAPSE for length (ventricular or wall), should reduce the part of biological variability due to body size (heart size). In the HUNT study, however, we found that both segmental strain by tissue Doppler (17), as well as by the linear method (MAPSE normalised for wall length) (18), compared to non-normalised MAPSE with relative standard deviations were similar for all three methods.
The finding that normalisation for LV length did not reduce biological variability, was perhaps a bit surprising, but is explainable as age is the source of the most variability (18)
Age | MAPSE (cm) | MAPSE vol(ml) | LVEDV(ml) | SV(ml) | EF(%) | MAPSE% of SV | Endocardial FS(%) | Outer FS (%) | Myocardial volume d (ml) |
Women | |||||||||
<40 | 1.73(0.20) | 56.5(9.9) | 111.6(21.6) | 76.3(16.4) | 68(6) | 75.4(11.9) | 36.6(6.1) | 14.1(3.3) | 87.0 (19) |
40-60 | 1.58(0.23) | 53.3(11.7) | 106.9(21.7) | 72.7(17.0) | 68(6) | 74.9(13.5) | 36.5(6.9) | 13.2(4.2) | 92.8 (19.6) |
>60 | 1.33(0.26) | 45.2(10.1) | 97.9(19.7) | 65.4(16.9) | 66(9) | 72.0(21.9) | 36.0(9.1) | 12.1(4.2) | 95.6 (18.9) |
Total | 1.58(0.26) | 52.9(11.5) | 106.8(21.8) | 72.6(17.3) | 68(6) | 74.6(14.9) | 36.4(7.1) | 13.3(4.0) | 91.4 (19.6) |
Men | |||||||||
<40 | 1.72(0.22) | 70.1(14.9) | 144.8(30.5) | 96.1(22.9) | 66(8) | 74.9(14.2) | 35.5(6.9) | 12.6(3.7) | 125.3 (23.6) |
40-60 | 1.58(0.22) | 65.1(14.2) | 138.1(31.1) | 92.2(23.8) | 67(8) | 72.8(14.8) | 35.8(7.4) | 12.2(3.8) | 129.7 (25.3) |
>60 | 1.45(0.21) | 60.3(14.7) | 126.3(33.7) | 84.1(25.7) | 66(8) | 74.9(19.0) | 36.0(8.0) | 11.8(4.4) | 128.2 (26.8) |
Total | 1.58(0.24) | 64.9(14.7) | 136.6(32.2) | 91.0(24.4) | 67(8) | 73.8(15.8) | 35.8(7.5) | 12.2(3.9) | 128.4 (25.3) |
All | 1.58(0.24) | 61.5(13.0) | 121.1(31.1) | 81.4(22.9) | 67(8) | 75.2(12.8) | 36.1(7.3) | 12.8(4.0) | 101.4 (27.9) |
Basically, the findings of the functional measures were:
In this study, SV calculated from EDV-ESV in the ellipsoid model was 81.4ml, while Mitral area x MAPSE was 61.5 ml, = 74.2% of total SV. Circumferential shortening due to OUTER circ. (diameter) shortening, was 12.8%, and must make up the rest, 25.8% of SV.
Endocardial FS and EF did not decline with age, as has been shown many times before, (33, 34, 35, 250, 252, 253). But as internal FS is stable, and external FS decline with age, there is no compensatory increase in transverse (short axis/circumferential) function to maintain EF, it's simply the simultaneous decline in LVEDV and SV, that maintains their ratio. And this is not due to decrease in short axis internal diameter, but to a decrease in length (19).
MAPSE declined with age as described before (18). But just as for EF, the simultaneous decline in MAPSE and SV, maintains their ration, so the percentage of the SV due to MAPSE doesn't decline, and thus there is no need for any increase in transverse function. In fact, there is a decrease in midwall and outer FS, so there is a decline in transverse function with age (7).
In addition, MAPSE was independently related to DBP (Age Beta = - 47%, DBP Beta -0.13, both p<0.001), but not to SBP. This means, that the (hypertrophic?) effect of hypertension, which was increasingly present with increasing age as discussed here, had an impact on MAPSE, but far less than age per see, and in this limited range, afterload (SBP) was not a factor.
MAPSE is nearly body size independent (18), while GLS is inversely related to body size (18) as explained here. The percentage of MAPSE contribution to SV, showed no correlation to BSA, whatsoever.
In HUNT 4 (245), based on the GE 2D strain speckle tracking method in 18 segments:
GLS was normally distributed.
Values were as follows (in numerical values):
< 40 years | 40 - 49 years | 50 - 59 years | 60 - 69 years | > 70 years | All | |
Women | ||||||
GLS (%) | 21.3 (1.8) | |||||
