Introduction
Much has been published in the last 20 years on the use of measurements of arterial stiffness in animal and human research studies. This summary statement was commissioned by the American Heart Association to address issues regarding the nomenclature, methodologies, utility, limitations, and gaps in knowledge in this rapidly evolving field. The following represents an executive version of the much larger Online Supplement, and is intended to give the reader a sense of why arterial stiffness is important, how it is measured, the situations in which it has been useful, its limitations and questions that remain to be addressed in this field. Throughout the document both “PWV” and variations such as “cfPWV” are used. The use of “PWV” (for pulse wave velocity) without modification is used in the general sense of arterial stiffness. The addition of lowercase modifiers like ‘cf’ (for carotid femoral, as in cfPWV) is used when speaking of specific segments of the arterial circulation.
The ability to measure arterial stiffness has been present for many years, but was invasive in the early times. The improvement in technologies to enable repeated, minimal-risk, reproducible measures of this aspect of circulatory physiology lead to its incorporation into longitudinal cohort studies spanning a variety of clinical populations including those at extreme cardiovascular risk (dialysis patients), those with co-morbidities like diabetes and hypertension, healthy elders, and general population surveys.
In the approximately three decades of clinical use of pulse wave velocity measures in humans we have learned much about the importance of this parameter. Pulse wave velocity has proven to have independent predictive utility when evaluated in conjunction with standard risk factors for death and cardiovascular disease. However, the field of arterial stiffness investigation, which has literally exploded over the last 20 years, has proliferated without logistical guidance for clinical and translational research investigators. This Summary Statement, commissioned by the American Heart association Council for High Blood Pressure Research represents an effort to provide such guidance, drawing on the expertise of experienced clinical and basic science investigators in Europe, Australia and the USA. Recommendations made in this statement are assumed to refer to the research aspect of arterial stiffness investigations, unless accompanied by language that emphasis clinical usage as well, and are based on the grid in Table 1.
Table 1.
Applying Classification of Recommendations and Level of Evidence
| SIZE OF TREATMENT EFFECT | |||||||
|---|---|---|---|---|---|---|---|
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CLASS I Benefit ⋙Risk Procedure/Treatment SHOULD be performed/administered |
CLASS IIa Benefit ≫ Risk Additional studies with focused objectives needed IT IS REASONABLE to perform procedure/administer treatment |
CLASS IIb Benefit ≥ Risk Additional studies with broad objectives needed; additional registry data would be helpful Procedure/Treatment MAY BE CONSIDERED |
CLASS III No Benefit or CLASS III Harm | ||||
| Procedure/Test | Treatment | ||||||
| COR III: No benefit | Not Helpful | No Proven benefit | |||||
| COR III: Harm | Excess Cost w/o Benefit or Harmful | Harmful to Patients | |||||
| ESTIMATE OF CERTAINTY (PRECISION) OF TREATMENT EFFECT |
LEVEL A Multiple populations evaluated* Data derived from multiple randomized clinical trials or meta-analyses |
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LEVEL B Limited populations evaluated* Data derived from a single randomized trial or nonrandomized studies |
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LEVEL C Very limited populations evaluated* Only consensus opinion of experts, case studies, or standard of care |
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| Suggested phrases for writing recommendations | should is recommended is indicated is useful/effective/beneficial |
is reasonable can be useful/effective/beneficial is probably recommended or indicated |
may/might considered may/might be reasonable usefulness/effectiveness is unknown/unclear/uncertain or not well established |
COR III: No Benefit | COR III: Harm | ||
| Comparative effectiveness phrases† | treatment/strategy A is recommended/indicated in preference to treatment B treatment A should be chosen over treatment B | treatment/strategy A is probably recommended/indicated in preference to treatment B it is reasonable to choose treatment A over treatment B | is not recommended is not indicated should not be performed/administered/other is not useful/beneficial/effective |
potentially harmful causes harm associated with excess morbidity/mortality should not be performed/administered/other |
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A recommendation with Level of Evidence B or C does not imply that the recommendation is weak. Many important clinical questions addressed in the guidelines do not lend themselves to clinical trials. Although randomized trials are unavailable, there may be a very clear clinical consensus that a particular test or therapy is useful or effective.
Data available from clinical trials or registries about the usefulness/efficacy in different subpopulations, such as sex, age, history of diabetes, history of prior myocardial infarction, history of heart failure, and prior aspirin use.
For comparative effectiveness recommendations (Class I and IIa; Level of Evidence A and B only), studies that support the use of comparator verbs should involve direct comparisons of the treatments or strategies being evaluated.
Section 1. What is arterial stiffness?
Recommendation
1.1 It is reasonable to measure arterial stiffness clinically by determining pulse wave velocity (Class IIa, Level of Evidence A) (1).
Arterial stiffness is a concept that refers to the material properties of the arterial wall, which in turn has functional consequences for the artery, since it affects the manner in which pressure, blood flow and arterial diameter change with each heartbeat. In addition to the passive mechanical properties of the load-bearing structures, arterial stiffness can be modulated by functional components related to cellular processes, where wall stiffness can be affected by endothelial function through modulation of smooth muscle tone or by altering the integrity of the extracellular matrix. As will be developed in this summary statement, stiffness is measured in different kinds of arteries (muscular, elastic), and in cross-section, longitudinally along the vessel, or in both directions. Often, arterial stiffness is assessed as the velocity of pulse wave travel in a defined segment such as the aorta. However, the research questions addressed by investigations of arterial stiffness are not restricted to this usage and stiffness has been measured in most named large arteries in humans (17). Arterial stiffness is also estimated by measuring pressure and/or diameter in a vessel and applying one or several of the now extensive formulae to the data to derive a value that reflects this inherent property of all arteries (18).
Surrogate measures of arterial stiffness – and what is not technically stiffness
Arterial stiffness is often determined by measuring the velocity of pulse wave travel in a segment of vessel (1). This is a valid measure, justified on the basis of equations such as Moens-Korteweg, or Bramwell Hill with which these measures agree (18). Other methods to measure arterial stiffness include the assessment of arterial compliance or distensibility, or measures of characteristic impedance (relating pressure changes to flow changes). When arterial geometry (size and wall thickness) is known, these can be used to compute the arterial wall elastic modulus, a direct expression of the stiffness of the wall. Confusion arises when measures such as systolic pressure augmentation, comparing the systolic pressure at the brachial artery with that at the proximal aorta, and sometimes reported as an augmentation index, are presented as “stiffness” parameters. Such measures are the result of several factors, including, but not limited to arterial stiffness (this is described further in the section on Wave Reflection) (19).
The arterial wall and stiffness
Arterial stiffness refers to the material properties of the arterial wall, which in turn affect the manner in which pressure, blood flow and arterial diameter change with each heartbeat. The pressure load of each heartbeat in large conduit arteries is borne mainly by the elastin and collagen components, with less contribution from smooth muscle in the muscular arteries. Due to the anatomical arrangement of the elastin and collagen fibers, elastin engages at low distension (hence at low pressure) and collagen at higher distension (and pressure) (20). The contribution of elastin and collagen to wall stiffness along the aorta varies as distance from the aortic valve increases, to optimize the reservoir function of the aorta.
Arterial stiffness is a major determinant of vascular impedance. Impedance relates the change in arterial pressure to the change in blood flow. Flow is determined by the presence of a pressure gradient. The relationships between time, pressure and flow are such that local wave velocity becomes a determinant of the instantaneous relationship between pressure and flow. For elastic conduits, the wave velocity is related to the stiffness of the wall, so changes in stiffness will modulate the pressure/flow relationships. The need to buffer each stroke volume, and to adapt to changes in flow, requires an optimal balance in the elastic and inelastic elements in the wall. Disease, aging, and other exposures typically reduce the elastic component and promote the inelastic (collagen) component such that arterial stiffness generally increases with age in most people.
Changes in arterial stiffness fall into passive and active categories. Passive categories relate to arterial wall fiber elements that are stretched and recoil with each heartbeat, and to heart rate (higher heart rates can be associated with increased arterial stiffness (16)). Active categories include endothelial function as it relates to nitric oxide and endothelin, as well as vascular smooth muscle where higher resting tone is associated with increase in arterial stiffness (21). Inflammation, oxidative stress, and turnover in the extracellular matrix of the vessel wall are additional active contributors to arterial stiffness (22). In addition, sympathetic tone and genetic polymorphisms appear to also regulate arterial stiffness in some vascular beds. The degree of the passive and active (functional) effects on wall stiffness depends on the type of artery, where a greater degree of functional effects would be manifest in more muscular arteries (e.g. carotid, iliac) compared to larger non-muscular conduit arteries (e.g. aorta) [callout i.ii here]
Section 2: Devices used to measure pulse wave velocity
Recommendations
2.1 You should non-invasively determine arterial stiffness by measuring carotid-femoral pulse wave velocity (Class I, Level of Evidence A) (2;3).
2.2 Pulse wave velocities measured in other vascular segments, such as ankle-brachial, or the determination of the cardiac-ankle vascular stiffness index, are useful in cardiovascular outcome predictions in Asian populations, but longitudinal studies in the US/Europe by these methods are lacking (Class I, Level of Evidence B) (4;5).
2.3 Single point estimates of pulse wave velocity are not recommended because there is a lack of evidence of cardiovascular outcome prediction in longitudinal studies. Measurement of pulse wave velocity in other arterial segments such as carotid-radial is not recommended because it does not predict outcomes (Class III, Level of Evidence B) (6).
Measurements of pulse wave velocity (PWV) are undertaken using several methodologies, some of which require sophisticated equipment (MRI) and software (some ultrasound & MRI devices). These fall into four categories:
Devices that use a probe or tonometer to record the pulse wave using a transducer
Devices using cuffs placed around the limbs or the neck that record pulse wave arrival oscillometrically
Ultrasound approaches
Magnetic resonance imaging (MRI)-based approaches
Devices using a probe or a tonometer to measure PWV
A number of devices based on this technology are available and have been extensively used in published research. Tonometry-based techniques (e.g. SphygmoCor ® device, AtCor Medical, Australia) uses a piezoelectric Millar tonometer which is placed at any two sites where a pulse is detectable. There is only one tonometer attached to the unit so PWV measurements require two sequential 10–20 second readings, gated to the ECG, to be taken. The average transit time is then derived using the R wave of the ECG as a reference point, and PWV calculated based on the inputted distance measurement. The SphygmoCor device has been used in Anglo-Cardiff Collaborative Study of arterial stiffness (23) and the Chronic Renal Insufficiency Cohort (CRIC) study of chronic kidney disease (24), as well as other cohorts and intervention studies. Newer versions of this device use a cuff and tonometer system to record simultaneous pressure waves (25). Published reproducibility of the PWV with the SphygmoCor, as judged by Bland-Altman plot analysis, is good (26).
Mechano-transducer based techniques (e.g. Complior ®, ALAM Medical, France) uses similar principles, but allows simultaneous measurement between sites using distension sensors. The Complior software provides an on-line pulse wave recording and automatic calculation of the PWV (27). This device has been used extensively in epidemiologic studies in Europe and provided much of the early outcome data relating PWV to CVD risk. The published reproducibility of the PWV with the Complior, as judged by Bland-Altman plot analysis, is good (28).
Other tonometry-based devices (e.g. PulsePen ®, DiaTecne, Italy) use an ECG signal and a hand held tonometer (similar to the SphygmoCor) to perform carotid-femoral PWV measures. The PulsePen has been used in the Predictive value of blood pressure and arterial stiffness in institutionalized very aged population (PARTAGE) study conducted in elderly patients in France and Italy (29). The reproducibility of the PulsePen, as judged by Bland-Altman plot analysis, is good (30).
Still other tonometry-based devices (e.g. those used by Cardiovascular Engineering, Inc., Norwood, MA) use a custom device to measure PWV using tonometric methods. The system uses the foot-to-foot measure of carotid and femoral pressure waveforms, with distance measures to the carotid artery site and femoral artery site calculated from the sternal notch. The ECG QRS complex is used as the timing onset point and the elapsed time to the carotid pressure waveform foot and the femoral pressure waveform foot is calculated and divided into the distance measurement. This system has been used in the Framingham (31) and Reykjavik Studies (32), as well as other cohorts and intervention trials. Reproducibility of the PWV by this method is, reportedly, good (Gary F. Mitchell, personal communication).
Devices using cuffs placed around the limbs or the neck that record pulse wave arrival oscillometrically
Oscillometric-based devices (e.g. VP1000 ®, Omron Healthcare, Japan) rely on four oscillometric cuffs – placed on both arms (brachial) and ankles to calculate brachial-ankle pulse wave velocity (baPWV). It also provides an ankle-brachial index (ratio of systolic pressure in the ankle compared with that of the brachial artery; a marker of peripheral arterial disease when this ratio is < 0.9). Newer models (e.g. VP2000) have additional probes that can be secured in place (with straps) that detect carotid (CAP) and femoral (FAP) pulses simultaneously (i.e. both probes capture the same pulse wave) by tonometry. ECG leads are attached, as is a phonocardiographic microphone (whether the measurements are being done by oscillometry or tonometry). The subject’s age, height and gender are entered into the software and the distance estimate is calculated using statistical norms (based on Japanese individuals). The Omron device has been used in prospective observational studies, mainly in Asia, and independently predicting loss of kidney function (33), cardiovascular disease (34), and all cause death (35). Published reproducibility of the PWV with the VP1000, as judged by Bland-Altman plot analysis, is good (36).
Cuff-based devices (e.g. Mobil-O-Graph ® (IEM, Stolberg, Germany) capture brachial blood pressure and brachial waveforms (causal and 24 hour), to estimate central aortic pressures, and to estimate carotid-femoral pulse wave velocity (37;38). The Mobil-O-Graph 24 h PWA ABPM device uses a proprietary algorithm to obtain conventional brachial blood pressure readings, after which the brachial cuff is inflated to the diastolic blood pressure level and held constant for approximately 10 seconds to record the pulse waves. Subsequently, central pressure curves are obtained using a transfer function. To estimate aortic PWV, several parameters from pulse wave analysis, along with wave separation analysis are combined in a proprietary mathematical model incorporating age, systolic pressure, and aortic characteristic impedance (39). The Mobil-O-Graph aortic PWV values have been validated by direct intra-arterial measurement in the catheterization laboratory (40). Reproducibility of the Mobil-O-Graph, as judged by Bland-Altman plot analysis, is good (41).
Some cuff-based devices (e.g. Vasera ®, Fukuda Denshi, Japan) use cuffs on all four limbs and gate the timing for the pulse wave arrival at the ankle relative to the heart using phonocardiography through a small microphone taped onto the chest (42). In addition to the cardio-ankle vascular index (‘CAVI’; which is derived from the cardio-ankle PWV) it also provides an ankle-brachial index. This device has been used mainly in Japan for longitudinal studies of dialysis patients (4) as well as in community studies of cognitive decline (43). Reproducibility of the Vasera, as judged by Bland-Altman plot analysis, is good (44).
Ultrasound approaches
Ultrasound can be used to assess vessel distension and derived stiffness indices, or flow waveforms to calculate PWV. Distension waveforms can be assessed using ultrasound transducers at a variety of locations, but often the carotid and femoral sites are employed. Although some parts of the aorta itself are assessable, measurements in the thoracic aorta are technically challenging. An average change in cross sectional area of a vessel can be derived from the distension waveform, using dedicated software (e.g. ARTLAB ®, ESAOTE, Italy). With a value for the pulse pressure, the operator can determine distension and compliance. Often brachial artery pressure is used rather than local pulse pressure, which may introduce inaccuracies, as may any delay between distension and BP assessment. Pulse wave speed (c) and other indices of elasticity – such as incremental elastic modulus (Einc) - can also be derived as discussed earlier. It is worth noting that most ultrasound systems and software produce a time average waveform, and mathematically this will yield different values for stiffness indices compared to calculating distension beat-by-beat and then averaging.
In addition, ultrasound is used to assess local (cross sectional) distensibility of vessels such as the carotid artery. B-mode ultrasound, video analysis and echo-tracking methodologies are common approaches used (45;46). The Online Supplement (Section 6) has more discussion of this aspect and device comparisons (Online Supplemental Table 6.4).
Doppler ultrasound may be used to record flow waveforms from accessible sites from which PWV can be estimated in a similar manner to PWV based on pressure waveforms. Waveforms may be recorded either sequentially with ECG gating, or simultaneously (47). Typically one ultrasound transducer is clamped to the left side of the neck to insonate the site of the left subclavian artery, or carotid artery, and the second transducer is secured on the abdomen insonating the abdominal aorta above the bifurcation. Distance is measured from the supra-sternal notch to the location of the second transducer. This can be challenging, because the angle of insonation makes it difficult to reliably determine where the abdominal aorta is being interrogated in most (obese) people. The foot of the flow wave from each of the recording sites is used, and the time elapsed in milliseconds is calculated. There is no set duration of recording, but averaging several beats (commonly 5–10 beats) is beneficial to increase the accuracy of the measurement (48). Identifying the foot of the flow wave can be more challenging than the foot of a pressure wave. However, such techniques have shown independent predictive value for cardiovascular outcomes, and death, in longitudinal studies of diabetics (48), the healthy elderly (49) and a general population (12). Published reproducibility of ultrasound-based PWV, as judged by Bland-Altman plot analyses, is good (50;51)
Magnetic resonance imaging (MRI)-based approaches
MRI can be applied in much the same way as ultrasound to determine distension-based indices or PWV. It has the advantage of being able to assess almost any vessel, and providing more accurate distance and area estimates –vessels can always be ‘cut’ in a perpendicular manner. However, these are offset against poorer time and spatial resolution and cost.
Phase contrast MRI (PC-MRI) can be used to acquire blood flow velocity maps along any given anatomical plane. When the gradient direction is applied exactly perpendicular to the cross-sectional vessel plane (“through-plane” velocity encoding), flow can be measured through the vessel cross section. Such an approach can be used to compute the time delay between the onset of flow in the ascending and descending thoracic aorta, which can be simultaneously interrogated in cross-section in a properly prescribed anatomic plane. Alternatively, the gradient direction can be prescribed in-plane with the vessel flow axis, allowing the acquisition of a velocity map along the length of the vessel. This approach allows the measurement of the spatiotemporal flow profile along the length of the vessel, thus allowing the computation of pulse wave velocity. This approach can be easily applied to the thoracic aorta in the “candy-cane” plane.
PC-MRI sequences require a user-defined velocity-encoding sensitivity (VENC), which should be as low as possible to minimize noise during the acquisition, yet higher than peak flow velocity in the region of interest to avoid aliasing. Although VENC should be tailored to individual measurements, a VENC of 130–150 cm/sec allows for an adequate interrogation of thoracic aortic flow in most cases. PC-MRI data are acquired over several cardiac cycles and consistent cardiac timing in each cycle is assumed. Adequate PC-MRI flow measurements require careful attention to technical details, including the recognition and minimization of sources of error such as phase-offset errors caused by in-homogeneities of the magnetic field environment (short-term eddy currents) (52;53), signal loss due to turbulent flow, partial volume averaging due to limited spatial resolution, signal misregistration due to in-plane movement of the aorta and pulsatile flow artifacts. The temporal resolution of PC-MRI flow measurements should be maximized, which requires data collection over multiple cardiac cycles. This is usually achieved by prolonged (several minutes) acquisitions during free breathing. Various alternative techniques have been proposed for fast, real-time assessments of PWV (54–57). More research is needed in regards to the optimal algorithm to measure the time delay between the foot of the flow waves using phase-contrast MRI.
A second approach to measure arterial stiffness with MRI involves the assessment of arterial distension, which can be paired with pressure measurements to obtain local arterial compliance and distensibility. Steady-state free precession techniques provide high contrast between the arterial lumen and arterial wall and allow for automatic segmentation of aortic lumen throughout the cardiac cycle. Such approaches can be used for the assessment of ascending aortic properties, as long as simultaneous (or quasi-simultaneous) central pressure recordings are performed. Unfortunately, tonometric arterial pressure recordings are difficult within the MRI suite, since available tonometry systems are not MRI-compatible. Good reproducibility of PWV by phase-contrast MRI has been reported, with intra-class correlation coefficients ~0.90 (58).
Many of the devices reviewed in this section can also be used to capture waveforms for central aortic pressure wave analysis. Section 4 in this executive summary, and Section 4 in the Online Supplement provide greater detail about this.
Irrespective of the approach used, it is critical to include an accurate measurement of blood pressure at the time of stiffness measurement because the mean arterial pressure is an important determinant of stiffness (see section 7 and recommendation 7.1). Reproducibility is generally good and most devices/approaches have been in use for at least a decade. Other approaches to measuring arterial stiffness are covered in the Online Supplement Section 2.
Section 3. Why is arterial stiffness important?
Recommendation
3.1 It is reasonable to measure arterial stiffness to provide incremental information beyond standard cardiovascular disease risk factors in the prediction of future cardiovascular disease events (Class IIa; Level of Evidence A) (3)
Arterial stiffness as a predictor of future cardiovascular risk
Stiffening of the central arteries has a number of adverse hemodynamic consequences, including a widening of pulse pressure, a fall in shear stress rate, and increased transmission of pulsatile flow into the microcirculation. These effects have a number of detrimental consequences that may, in part, explain mechanistically why stiffness is a predictor of risk. Numerous studies involving various disease-specific and community-based cohorts have demonstrated that higher carotid-femoral PWV (cfPWV) is associated with increased risk for a first or recurrent major cardiovascular disease event (2;3). Consideration of cfPWV substantively reclassifies risk in individuals at intermediate risk for CVD, suggesting that consideration of cfPWV provides novel and clinically relevant information beyond that provided by standard risk factors (3;31). In addition, small studies have demonstrated that persistent elevation of cfPWV during treatment for hypertension or CVD is associated with high risk for an adverse outcome in those with established disease (59;60). The added benefit of cfPWV in risk prediction models may be a manifestation of the relatively modest relation between cfPWV and standard risk factors other than age and blood pressure (61). In a recent individual participant meta-analysis, higher cfPWV was shown to be associated with increased risk for coronary heart disease, stroke and composite cardiovascular events. Importantly, relative risk was strongest in younger individuals, where an opportunity exists for early identification, lifestyle modification and possible mitigation or prevention of further potentially irreversible deterioration of aortic structure and function (3).
Hypertension
The association between arterial stiffness and hypertension is well established (62–66). There is a widely held belief that increased aortic stiffness in hypertensive individuals is largely a manifestation of longstanding hypertension-related damage that stiffens the large arteries. A recent analysis from the Framingham Heart Study found that higher arterial stiffness, as assessed by cfPWV, was associated with blood pressure progression and incident hypertension 7 years later (62). However, higher blood pressure at an initial exam was not associated with progressive aortic stiffening, suggesting that aortic stiffness is a cause rather than a consequence of hypertension in middle-aged and older individuals. These results and several additional studies provide strong evidence in support of the hypothesis that arterial stiffness represents a cause rather than a consequence of hypertension and underscore the importance of better defining the pathogenesis of aortic stiffening (63–66).
High aortic stiffness is associated with increased blood pressure lability (67–69). A stiffened vasculature is less able to buffer short-term alterations in flow. Increased aortic stiffness is also associated with impaired baroreceptor sensitivity (67;70–72). Together, these limitations may result in potentially marked alterations in blood pressure as cardiac output changes during normal daily activities, such as changes in posture and physical exertion (73).
Cardiac disease
Excessive arterial stiffness represents a compound insult on the heart. Aortic stiffening increases left ventricular systolic load, which contributes to ventricular remodeling and reduced mechanical efficiency. This leads to an increase in myocardial oxygen demand (74), compounded by a reduction in diastolic coronary perfusion as pulse pressure widens and diastolic blood pressure falls with aortic stiffening (75). Arterial stiffening may be associated with impaired measures of left ventricular diastolic function (76;77), which may increase cardiac filling pressure and further limits coronary perfusion. Finally, arterial stiffness is associated with atherosclerosis (78–81), which may further impair ventricular perfusion, possibly leading to catastrophic reductions in ventricular function during ischemia (75).
Arterial stiffness is associated with diastolic dysfunction and diastolic heart failure due to direct effects of abnormal load and loading sequence on myocyte contraction and relaxation and indirectly through ventricular hypertrophy (77;82–86). Diastolic dysfunction increases filling pressures and thus may increase load on the atria, which will contribute to atrial hypertrophy and fibrosis and ultimately to atrial fibrillation (87). Arterial stiffness is independently associated with an increased risk of heart failure (88) and is increased in patients with established heart failure whether left ventricular function is preserved or impaired (89–91).
Peripheral vascular function
Arterial stiffness (arteriosclerosis) is associated with atherosclerosis, although the association is not a strong one and the two processes should be viewed as distinct pathophysiological entities. Aortic stiffening may increase the risk for development of atherosclerosis as a result of atherogenic hemodynamic stresses associated with a stiffened aorta, including increased pressure pulsatility and abnormal flow patterns in large arteries, with high flow and shear stress during systole and stasis, or flow reversal, during diastole (92). Arteriosclerosis also has important implications for structure and function of the microcirculation.
Aortic stiffening leads to loss of normal impedance mismatch between the normally compliant aorta and stiff muscular arteries. Loss of impedance mismatch reduces the amount of wave reflection at the interface between aorta and proximal branch vessels and therefore increases transmission of excessive pulsatile energy into the periphery where it may cause damage (32;93;94). Increased aortic stiffness and excessive pressure pulsatility are associated with increased resting microvascular resistance and markedly impaired post-ischemic reactive hyperemia in the forearm (95). Resistance vessel remodeling, as assessed by the media-lumen ratio, is more closely related to pulse pressure than mean pressure, suggesting that anatomical constraints may contribute to limited reactivity in remodeled vascular beds (96–99). Indeed, a recent study demonstrated a significant relationship between aortic PWV and media-lumen ratio of small resistance arteries in a cohort of hypertensive patients after adjustment for age and blood pressure (100). Dynamic tone in small arteries is also affected by pressure pulsatility (101–104). As a result, vascular resistance in autoregulated organs such as the kidney and brain may depend on pulse pressure as well as mean arterial pressure. If resistance vessel tone increases in response to pulse pressure at a constant level of mean pressure, flow will fall as resistance increases. Hence, alterations in the relation between mean and pulse pressure could lead to dissociation between mean pressure and resistance and interfere with autoregulation of flow. Beyond midlife, pulse pressure increases rapidly as mean pressure remains constant or falls, potentially putting autoregulated organs at risk for relative ischemia.
Central Nervous System
High flow organs such as the brain and eye are particularly sensitive to excessive pressure and flow pulsatility (105). High local blood flow is associated with low microvascular impedance, which facilitates penetration of excessive pulsatile energy into the microvascular bed (32). This may contribute to repeated episodes of microvascular ischemia and tissue damage, and manifests as white matter hyperintensities, clinically unrecognized focal brain infarcts, and tissue atrophy, each of which contributes to cognitive impairment and frank dementia.
Aortic stiffening is also associated with increased risk for large vessel stroke, which may be ischemic or hemorrhagic (106;107). This may be mediated through atherosclerosis, with increased stiffness contributing both to atherogenesis and risk for plaque rupture (108), through atrial enlargement and fibrosis which can trigger atrial fibrillation, providing a cardiac source for embolus (87), or through diastolic flow reversal in the aorta, which could disrupt and redirect plaque from the distal arch into the carotid circulation (109). Excessive pressure pulsatility can also predispose to large artery dissection or rupture of intracranial aneurysms, leading to hemorrhagic stroke. In addition, increased aortic stiffness is associated with blood pressure lability, which is a risk factor for incident stroke (110).
Arterial stiffness is also associated with impaired cognitive function in selected (111–115) and community-based samples (32;116–120). In light of the generalized insult on the brain vasculature that occurs, it is perhaps not surprising that aortic stiffness is associated with a broad spectrum of cognitive sequelae, and has been established as a risk factor for both vascular and Alzheimer-type dementia (121).
Kidney disease
Like the brain, the kidneys are low impedance organs that are exposed to obligate high flow throughout the day. In addition, the unique structure of the microvasculature in the kidney, with resistance vessels on either side of the glomerulus, markedly increases pressure in the glomerulus nearly to aortic levels. In the presence of increased aortic stiffness, the microvasculature of the kidney is exposed to excessive pressure and flow pulsatility that can damage the glomerulus, leading to proteinuria and loss of function (122;123). Recently increased renal pulsatility has been correlated with CV and renal outcomes (124). Numerous studies have demonstrated modest but robust associations between increased pulse pressure or PWV and reduced glomerular filtration rate (GFR) or proteinuria (125–131). However, relations between estimated GFR and stiffness measures are less robust in some studies after adjusting for potential confounders. In a study that measured GFR directly, higher PP was associated with reduced measured GFR (132). Importantly, PP was not related to GFR estimated from serum creatinine in that study, indicating that relations between PP and estimated GFR may be obscured in older individuals, where loss of muscle mass may reduce accuracy of creatinine-based GFR estimating equations (133–135). Given that the prevalence of abnormal aortic stiffness is heavily age-dependent, the burden of stiffness-related kidney damage may be underestimated when estimated GFR is used as a surrogate for kidney function.
Thresholds and normative values for risk assessment
cfPWV was included in the 2007 ESH/ESC guidelines for the management of hypertension (136), where a fixed cutoff of 12m/s was proposed, indicating subclinical organ damage. This was modified by a recent expert consensus, which took into consideration a new distance calculation methodology, and recommended a new 10 m/s threshold (derived by multiplying 12 m/s by 0.8 and then rounding up) (137). Although attractive because of the simplistic approach, risk estimation based on fixed thresholds has several limitations, not least of which are the relatively continuous relationship between risk and cfPWV, and the failure to consider factors such as transient elevation of MAP, which may confound cfPWV values because of nonlinear stiffness of the aortic wall.
A single threshold also fails to take into consideration the dominant effect that age has on PWV. A cfPWV value of 12.1 m/s may convey different prognostic information in an 80 year-old vs. a 25year-old person. Variability of cfPWV with age prompted an interest in attempting to establish reference values for various segments of the population (137;138). The European Network for Non-Invasive Investigation of Large Arteries assembled the Reference Values for Arterial Stiffness’ Collaboration whose task was to generate reference and normative values for cfPWV. The cohort included 11,092 individuals, who yielded reference values of cfPWV stratified by age groups (in years: <30, 30–39, 40–49, 50–59, 60–69 and >70). In addition, from the subset of individuals who had optimal or normal blood pressure and no additional cardiovascular risk factors, normative values for cfPWV were also generated according to age groups (139). However, it should be emphasized that these normative and reference values are predominantly applicable to measurements performed using the aforementioned methodologies.
Despite the attractiveness of age-relative normative thresholds, it is important to recognize an age-related increase in cfPWV should not necessarily be viewed as inevitable or indeed a normal part of the ageing process. Although cfPWV increases exponentially with aging in most populations, it appears to rise much less rapidly in truly rural or indigenous populations (140;141), as Trusswell reported for blood pressure in the 1970s (142). The observation that cfPWV increases more modestly with age in lower risk individuals suggests that a major part of age-related stiffening is pathological and as such, it may not be appropriate to use age-specific thresholds for risk estimation.
Section 4: Arterial stiffness, wave reflections and LV afterload
Recommendations
4.1 We recommend that you quantify both time-resolved central pressure and central aortic flow when assessing LV afterload, as either an exposure for a cardiovascular outcome, or a target for intervention (Class I; Level of Evidence C).
4.2 We recommend the use of pressure-flow analyses, which are considered the gold standard assessment, to determine left ventricular (LV) afterload (Class I, Level of Evidence A) (7;8).
4.3 Effective arterial elastance (Ea) should not be used as an index of pulsatile LV afterload or arterial stiffness, since it represents a poor index of pulsatile load and is not significantly influenced by arterial stiffness (Class III, Level of Evidence B) (9;10).
4.4 We recommend the use of wave separation analysis, as opposed to aortic augmentation index, when investigations are specifically focused on the role of wave reflection as either an exposure for a cardiovascular outcome or as a target for intervention (Class I, Level of Evidence B) (11–13).
The mechanical “afterload” imposed by the systemic circulation to the pumping left ventricle (LV) is the aortic input impedance, and is an important determinant of normal cardiovascular function as well as a key pathophysiologic factor in various cardiac and vascular disease states. In the presence of a normal aortic valve, LV afterload is largely determined by the elastic properties (arterial stiffness), arteriolar caliber and wave reflection characteristics of the arterial tree (“arterial load”) (8). Arterial load is complex, time-varying, and cannot be characterized by a single number or index. LV afterload is composed of a “steady” and a “pulsatile” component and can be described by the following indices: (1) Systemic vascular resistance (SVR); (2) Aortic characteristic impedance; (3) Total arterial compliance; (4) Wave reflection amplitude; (5) Reflected wave transit time.
SVR, the steady component of LV afterload, is largely determined by arteriolar caliber and number. Pulsatile load, in contrast is determined by the hemodynamic function of conduit arteries, which in turn depends on their geometry and wall stiffness. Although brachial arterial pressure (systolic, diastolic and pulse) is often used as a surrogate of arterial function and LV afterload in clinical practice, LV afterload cannot be fully described in terms of peripheral pressure alone and needs to be assessed in the frequency domain from central aortic pulsatile pressure-flow relations (143;144), or estimated in the time domain from the aortic pulsatile pressure alone (18).
Furthermore, it should be recognized that: (1) afterload affects, in a reciprocal fashion, the pressure and flow waves generated by the LV; (2) pressure and flow waves are not only dependent on load, but are also strongly influenced by LV structure and function.
Flow can be measured invasively using a flow wire, or non-invasively with MRI or with pulsed wave Doppler echocardiography interrogating the LV outflow tract (LVOT). Central aortic pressure can be measured invasively with a pressure sensing catheter or wire, or via radial arterial tonometry and a general transfer function (GTF) which synthesizes a central aortic pressure waveform (18), or by carotid arterial tonometry. For non-invasive assessments, calibration of central pressure waveforms should be performed using peripheral diastolic and mean arterial pressures, which (in contrast to systolic pressure) remain relatively constant throughout the arterial tree (7). To obtain central aortic pressure waveforms, calibration of the radial artery waveform should be performed using peripheral systolic and diastolic pressures (1;18).
An increase in the pulsatile component of afterload causes an undesirable mismatch between the LV and the arterial system, increasing myocardial oxygen demand and decreasing cardiac efficiency (74;145). These changes in ventricular/vascular coupling promote the development of LV hypertrophy (LVH) and often lead to both systolic and diastolic myocardial dysfunction (see below) (77;146–148). In health, there is an increase (or amplification) in the pulse pressure as the pulse wave travels from the proximal aorta to the periphery. Increasing aortic wave reflection amplitude increases aortic systolic pressure and decreases the gap between central and peripheral pulse pressure and dampens (or reduces) this amplification. Decreasing wave reflection amplitude with antihypertensive therapy and/or exercise conditioning increases the gap (and amplification) and reduces target organ damage (149). Conversely, a reduction in pulse pressure amplification is associated with overt target organ damage and independently predicts future cardiovascular mortality (150;151).
Thus, pulse pressure amplification has been proposed as a potential mechanical biomarker of cardiovascular risk and global arterial function. Due to systemic changes in arterial stiffness and wave reflections coupled with changes in heart rate, brachial BP is not an accurate predictor of LV load and central hemodynamic burden. Moreover, the beneficial reduction in ascending aortic systolic and pulse pressures with various therapeutic approaches is often underestimated by cuff measurements of brachial artery pressure (18;152).
Once measures of central aortic pressure and flow are obtained, these can be modeled to assess steady, and pulsatile LV afterload and the amplitude and timing of wave reflections. An important relationship in the aorta is the pressure adaptation to pulsatile flow. When there is no influence on this relationship from wave reflections, as occurs early in early systole, pressure and flow wave forms look very similar. The relationship of aortic pressure and flow in the absence of wave reflections is called the characteristic impedance and is typically depicted as Zc (or Zo). An illustration of this relationship is shown in the Online Supplement (Figure 4.3). After arrival of the reflected wave in the central aorta, the pressure and flow waveforms diverge, because the reflected wave increases systolic pressure and reduces flow during deceleration. The degree of this divergence is associated with the local Zc and reflection site distance (153–155). This principle is used in linear wave separation analysis, which decomposes pressure and flow waveforms into their forward (incident) and backward (reflected) components. Reflection magnitude (RM) is expressed as the ratio of the amplitudes of reflected/forward pressure waves (155), while reflection (or augmentation) index (AIx) is the ratio of the amplitude of the reflected wave and central aortic pulse pressure. These two variables (RM and AIx) are similar and are measures of wave reflection strength (or intensity). Reflected pressure waves arriving at the proximal aorta increase the late systolic load of the left ventricle, thus altering the loading sequence. Increased wave reflection amplitude and/or a left ventricular loading sequence characterized by late systolic load have been shown to cause myocardial hypertrophy (146;156;157), myocardial fibrosis (156), systolic and diastolic myocardial dysfunction (82;84;158–164) and to strongly predict an increased risk of future heart failure (11;163). Increased wave reflections have also been shown to predict all-cause 15-year mortality (12).
Since invasive recordings of central aortic pressure and flow waves and pulse wave analysis can only be made in a select number of patients in the catheterization laboratory, techniques have been developed recently that enable the non-invasive determination of the above variables (165;166) in large cohorts with similar results (145;167–172)). Some studies use the carotid artery wave as a surrogate for the central aortic pressure wave while others derive it from the radial artery wave using a GTF. Briefly, radial artery pressure waves are recorded at the wrist, using applanation tonometry with a high-fidelity micromanometer. After 20 sequential waveforms are acquired and ensemble averaged, a validated GTF is used to synthesize the central aortic pressure wave non-invasively. To obtain the GTF, computer software performs a Fourier series representation of the radial artery waveform into harmonic components of amplitude and phase angle. These harmonics are then adjusted using data obtained from previous invasively measured aortic pressure waves to obtain the non-invasive synthesized central aortic pressure wave (3). Two visible demarcations usually occur on the initial upstroke of the central aortic pressure wave in middle-aged and older individuals; the first shoulder and the inflection point. These demarcation points occur at an earlier age in patients with hypertension. The first (or early) shoulder is generated by LV ejection and occurs at peak blood flow velocity while the inflection point occurs later and denotes the initial upstroke of the reflected pressure wave; this wave represents the second (or mid-to-late) systolic shoulder (18;173–177). The first shoulder is an estimate of forward traveling wave amplitude while the second shoulder is an estimate of reflected wave amplitude. The characteristics of the reflected wave depends upon the physical properties (stiffness, taper and branching) of the entire arterial tree (elastic plus muscular arteries and arterioles), PWV, the round-trip travel time of the wave from the heart to the periphery and back, and the distance to the major “effective” reflecting site in the lower body (18;173–177).
Effective arterial elastance (Ea), computed as the ratio of end-systolic pressure to stroke volume, was proposed as a lumped parameter of resistive and pulsatile LV afterload (178) and is increasingly used due to the simplicity of its computation. However, Ea is almost entirely determined by the product of heart rate (a cardiac property) and SVR (179;180), and (despite its name) does not reflect or characterize pulsatile LV afterload (9;10). Ea does not represent a physical elastance (or compliance) and is not related to arterial stiffness. Therefore, it should not be interpreted or used to measure pulsatile afterload or arterial stiffness.
Interventions that reduce arterial stiffness and wave reflections, the primary cause of elevated systolic blood pressure and LVH, include drugs prescribed for the treatment of hypertension, and heart failure. These drugs are usually categorized as vasodilators, aldosterone blockers, β-blockers and diuretics. Different cardiovascular drugs have different effects on arterial properties (structure and function) and wave reflection characteristics (167;181–183). In most countries, thiazide diuretics are the cheapest antihypertensive drugs available and they are the recommended first-line treatment for hypertension in the US (JNC 7) (184). Diuretics and pure β-blockers decrease blood pressure but have little, if any, direct (active) effect on arterial properties and wave reflection characteristics. Selective and nonselective aldosterone blockers attenuate cfPWV and AIx (185;186) in select patient groups by increasing nitric oxide (NO) bioactivity and improving endothelial vasodilator dysfunction (187). Vasodilating drugs, such as hydralazine and dipyridamole, primarily increase arteriolar caliber and therefore decrease peripheral resistance and mean arterial pressure via their action on arteriolar smooth-muscle cells with little effect on aortic wave reflections (188). Nitrates primarily relax smooth muscle cells in large conduit muscular arteries and therefore decrease arterial stiffness, aortic wave reflection amplitude and duration and reduce central systolic and PP with little change in brachial cuff systolic and PP (189–191). ACE inhibitors, ARBs and CCBs are the most commonly used vasodilator drugs. These drugs appear to have little direct effect on stiffness of elastic arteries like the aorta independent of blood pressure reduction (18;183), although some studies question this (192–194). A recent meta-analysis observed that ACE inhibitor therapy improved the stiffening of arteries as reflected by PWV and reduced arterial wave reflections as assessed by augmentation index (AIx) when compared with placebo (195). Beta-blockers appear to show less benefit on central aortic pressure compared with ACE inhibition (e.g. in CAFÉ (196)), but less is known about newer beta-blockers that feature either concurrent alpha blockade (carvedilol) or nitric oxide stimulation (nebivolol).
Several non-pharmacological interventions reduce arterial stiffness and wave reflections, including aerobic exercise training (197;198), dietary changes (including weight loss and salt reduction) (199–202), passive vibration (203) and enhanced external counter-pulsation (EECP) treatment (204). For maximum cardiovascular benefits these interventions must be initially introduced acutely and continued over an extended period of time. While the effects of exercise on arterial stiffness and wave reflections have been studied for more than half a century (205), many aspects remain unclear. It appears that the effects depend on the type (aerobic or resistance), on the intensity, and on the duration of exercise [acute or chronic (endurance, training, conditioning)]. Multiple studies attest to the benefits of regular aerobic physical exercise in advanced age, hypertension, diabetes, coronary artery disease and heart failure and to the improvement in oxygen extraction from blood, and in cardiovascular function that occur with exercise training. Cross-sectional studies of aerobic exercise trained individuals are conflicting and have reported both reduced pressure from wave reflections (198;206;207) and increased pressure from wave reflections (208;209). These differences in wave reflection characteristics and central aortic pressure may be linked to lower heart rates in the endurance trained subjects. The increase in pressure is probably due to an increased in the first systolic shoulder resulting from an increase in peak aortic blood flow. Longitudinal exercise training studies are similarly somewhat conflicting and have noted improvements in pressure from wave reflections (197;210) or no change (211). Although endurance exercise training has been shown to reduce arterial stiffness and improve peripheral vascular tone and endothelial function, exercise training-mediated reductions in heart rate (212) and improvements in LV contractility (213) likely represent equipoise in their potential to detect a reduction in pressure from wave reflections consistently across studies. There is no doubt that weight loss and regular exercise lowers LV afterload (static and dynamic components) and heart rate, enhances quality of life and reduces morbidity and mortality from cardiovascular events (214). In a recent review of the effects of diet and exercise on arterial stiffness in patients with elevated cardio-metabolic risk from hypertension, signs of atherosclerosis, or kidney disease, Sacre and colleagues noted that these non-drug interventions can improve arterial stiffness by several mechanisms (215). Aerobic exercise may do so through improving vascular smooth muscle cell relaxation through increased nitric oxide bioavailability, and reductions in oxidant stress and inflammation. Among dietary approaches, reduced sodium intake has so far shown that while it is associated with reductions in pulse wave velocity, these seem due largely to the changes in blood pressure that occur (though others have found a reduction in PWV independent of blood pressure changes (216)). They (i.e. Sacre et al) also noted that increased sodium intake, and caffeine supplements, tended to promote arterial stiffness. People who exercise regularly are more likely than those who do not to control their weight and to control other risk factors for coronary and other vascular diseases. In older individuals, one year of exercise training was found to significantly improve physical fitness and lifetime risk for cardiovascular disease without affecting endothelial function or arterial stiffness (217).
Acute resistance exercise imposes a very different stress on the CV system than aerobic exercise. While aerobic exercise induces a volume load on the heart and other organs, resistance exercise imposes a pressure load. Acute resistance exercise increases pressure from wave reflections and unlike aerobic exercise, resistance exercise increases aortic stiffness and reduces pulse pressure amplification (218). The effect of habitual resistance exercise training on central aortic stiffness and pressure from wave reflections remains controversial. A recent meta-analysis concluded that high intensity resistance exercise training is associated with increases in central aortic stiffness in those with lower baseline stiffness values (219). Resistance exercise training was initially shown to increase pressure from wave reflections (220) with subsequent studies noting no effect (212;221–225).
Other aspects of ventricular-vascular coupling, including myocardial wall stress, are covered in the Online Supplement material Section 4.
Section 5. Arterial Stiffness in Children
5.1 Devices measuring stiffness in children should be validated in children (Class I, Level of Evidence C)
The participants of the major longitudinal studies of CV risk factors in children are too young to provide data linking CV risk factor levels measured in childhood to hard CV events in adulthood (226).
However, there are correlations between known adult CV risk factors, high-risk conditions, such as chronic kidney disease (CKD) and diabetes mellitus (DM), and novel risk factors with intermediate non-invasive measures of vascular health, which are linked to hard events in adults. In this section, we will discuss the current evidence, with reference to the previous AHA manuscript on non-invasive measures in children (227).
Arterial stiffness, CV risk factors, and high risk disease states in pediatrics
There are now sufficient data from studies such as the Bogalusa Heart Study, to link CV risk factors measured in youth such as blood pressure (BP) directly to estimate PWV in adulthood (228). The CV Risk in Young Finns study has also demonstrated higher adult PWV with clustering of risk factors in youth, such as in the metabolic syndrome (229). Conversely, clustering of advantageous risk factors (fruit and vegetable consumption) is associated with a lower PWV as an adult (230). Low birth weight was associated with higher PWV in adulthood in one study that examined baPWV (231), but was not found in a study that examined cfPWV (232). These differences highlight the importance of standardization of measurements, and that indices of stiffness are not always interchangeable as they may convey different predictive values.
These observations led to interest in delineating the determinants of PWV in healthy children and adolescents. Two studies evaluated gender differences in PWV. One study found higher carotid-femoral (cf) and femoral-dorsalis pedis (fd) PWV in females before puberty with the difference for cfPWV disappearing after maturation whereas fdPWV was higher in males post-puberty (233). Using baPWV, higher values were found in males regardless of maturation level (234).
Traditional CV risk factors have been found to influence PWV in youth. Children with elevated LDL-cholesterol had significantly higher PWV compared to controls (4.72 ± 0.72 m/sec vs 3.66 ± 0.55 m/sec) (235) and PWV increases across tertiles of TG/HDL ratio, a lipid parameter that reflects burden of small dense LDL particles (236). Higher PWV compared to controls was found in adolescents with a family history of hypertension (237;238), pre-hypertension (239–241), and sustained hypertension (239;240). Other CV risk factors such as psycho-social stress (242–244), smoking (245), low physical fitness (246;247), or physical inactivity (248;249), and low dairy intake (250) have also been related to higher PWV in pediatric patients. However, the studies vary considerably in adjustments for confounding factors such as MAP, heart rate and age, making interpretation of potential causality difficult.
Many data are also available to examine the relationship between obesity and PWV in the young. Two large studies with over 600 subjects each demonstrated higher PWV in obese adolescents as compared to their lean counterparts (251) and the effect of obesity was independent of other CV risk factors (252). Obesity-related metabolic syndrome clustering was also shown to result in higher PWV (253). However, insulin resistance appears to play an independent role only for baPWV (254;255) and not cfPWV (256).
Since CV risk factors influence PWV, it is not surprising that higher PWV is found in children and adolescents with high-risk conditions. Youth with type 2 DM have higher PWV than both their lean and obese counterparts (252). Surprisingly, subjects with type 2 DM have higher PWV than those with type 1 DM despite a shorter duration of disease (257). In a study of 535 subjects with type 1 DM and 60 with type 2 DM, it was found that the higher PWV in type 2 DM was largely explained by increased central adiposity and higher BP (257).
Pediatric patients with renal disease also demonstrate increased arterial stiffness, a potential mechanism for the observed increased in CV events in adults with kidney disease (258). Children on dialysis have higher PWV than less severely affected patients (259) and controls (260). Unfortunately, these adverse vascular changes may not normalize after renal transplantation (261–263). However, children with glomerulonephritis and increased PWV did see normalization with recovery (264). For this reason, there is hope that treatment of inflammatory vasculitis such as seen in HIV-infection (265), polyarteritis nodosa (266), and Kawasaki’s disease (267), may result in reduction of PWV although these types of long-term interventional studies have not been carried out to date.
A number of studies have evaluated PWV in children with congenital heart disease. Not surprisingly, PWV is higher in pediatric patients after cardiac transplant (268). Increased PWV has also been demonstrated after repair of tetralogy of Fallot, which is hypothesized to be a risk factor for progressive aortic root dilation in these patients (269–271) and in youth after arterial switch operation for transposition of the great vessels (272). The largest amount of work has been done in patients after repair of coarctation of the aorta, due to heightened concern for the role of arterial stiffness, manifest as increased PWV, in late complications such as hypertension (273–276) and premature CV disease (277). Data on other inherited disorders associated with increased arterial stiffness are less clear. One study of Marfan’s syndrome patients found higher PWV compared to controls (176), while another small study of Marfan’s (N= 10 cases and 10 controls) (278) and one of youth with neurofibromatosis type 1 (279), found no differences. Clearly, larger studies of PWV in pediatric patients with these high-risk conditions should be conducted.
The use of these non-invasive intermediate endpoints to better risk-stratify youth is essential as data linking childhood measures of CV risk factors to hard CV events in adults are lacking. Further studies correlating risk factors to vascular damage to target organ damage like LVH will provide evidence to pediatric practitioners faced with the challenge of implementing aggressive drug therapy in high risk children. Assessing PWV in ‘normal’ children may also provide an ideal platform to identify novel mechanisms driving stiffness as the influence of traditional cardiovascular risk factors and atherosclerosis per se will be minimized.
Developmental changes in arterial function in childhood
Many investigators have found an increase in arterial stiffness from childhood to adolescence (234;280–282), including large and small artery compliance (283). Using MRI, Voges found a decrease in descending aorta distensibility and increase in PWV starting at age 2.3 years (284). It appears this must relate to changes in the vessel wall since vascular compliance is determined by both vessel size and distensibility of the wall and the MRI study demonstrated a steady increase in cross sectional area of the descending aorta (with a slight plateauing after 15 years of age (284). Similarly, Senzaki found that although arterial compliance increased from birth to 20 years, once normalized for BSA to control for differences in arterial size, there was an overall decline over this period of time although the rate of change was not constant, with the most rapid decline in compliance during periods of most rapid growth from 3 to 7 years of age (285). Whether there are gender-related differences in developmental changes in arterial stiffness is less clear as Ahimastos found lower systemic arterial compliance and PWV in pre-pubertal girls compared to boys with no difference seen post-puberty (233), Fischer found sex differences in PWV both pre- and post-puberty (281), and Voges found no difference (284). Clearly more studies defining normal levels for arterial function parameters and better data outlining the determinants of increased stiffness across the pediatric age groups are needed. Other vascular measures such as arterial distensibility, Aortic Augmentation Index (AIx), ambulatory arterial stiffness index, normal values in youth, and technical considerations for measurement in children are also discussed in the Online Supplement Section 5.
Section 6. Validation of Arterial Stiffness devices
Recommendation
6.1 You should determine the distance for the carotid-femoral PWV by subtracting the suprasternal-notch to carotid site distance from the suprasternal-notch to femoral site distance, or by multiplying the total directly measured distance by 0.8 (Class I, Level of Evidence B) (14).
6.2 We recommend that you perform validation studies against invasive measurements; where this is not possible they should be validated against a non-invasive device that has been used in prospective trials showing an independent prognostic value of PWV [Table 2] ( Class I, Level of Evidence C).
Table 2.
[Section 6]: Recommendations for grading comparisons of devices/procedures for measuring PWV with a gold standard device
| Excellent accuracy: mean difference ≤ 0.5 m/sec and SD ≤ 0.8 m/sec |
| Acceptable accuracy: mean difference < 1 m/sec and SD ≤ 1.5 m/sec |
| Poor accuracy: mean difference ≥ 1 m/sec or SD > 1.5 m/sec |
In this section we review the standards by which measurement methods of PWV are validated, discussing several methodologies for non-invasive PWV estimation.
Invasive aortic PWV
This measurement has the advantage of being a simple, straightforward, precise, reproducible technique (measuring transit time (TT) simultaneously or ECG triggered and travel distance (TD) between two measurement sites (14). Of note, pressure waves measured at different points in the aorta travel only in one direction along the aorta, yielding a physiologically correct measurement. However, true invasive aortic PWV has been reported rarely and for obvious reasons only in patients scheduled for coronary angiography (14;286–291). So far, one study investigated its relationship to clinical outcomes (288).
Magnetic resonance imaging based aortic PWV
With this technique, TD can be measured very accurately due to precise 3 dimensional imaging approaches. TT can be estimated, using dedicated sequences to derive flow signals. Flow signals as measured travel along the aorta in only one direction along a single path, yielding a physiologically correct measurement. However, the temporal resolution for TT assessment is somewhat lower in comparison to the other techniques, although this has been improved recently (292). The reproducibility and the accuracy with respect to invasive measurements may also depend on the methods used to determine TT (293), and to date there is no consensus on the best method to be used. Finally, there are no published studies relating MRI-based aortic PWV with cardiovascular endpoints.
Simultaneous non-invasive acquisition of pressure waves at the carotid and femoral artery
There are no studies showing the superiority of simultaneous measurements as opposed to sequential (ECG-triggered) recordings. When the sequential recordings are made a short time apart, heart rate variability or the change in the isovolumic period probably have no or only minor effects on measured TTs (45).
Can dedicated devices for the measurement of cfPWV be recommended as a non-invasive gold-standard?
Validation studies using invasive aortic PWV as reference are limited to patients undergoing cardiac catheterization on clinical indications, thus limiting such studies to a relatively small group of patients. When MRI-based aortic PWV is considered as reference, the dedicated MRI environment often will preclude simultaneous measurements (the same is true for invasive aortic PWV). In addition, some questions with respect to temporal resolution remain to be solved. For these reasons, it seems reasonable to perform validation studies against dedicated devices that have been used widely in prospective trials showing an independent prognostic value of cfPWV (Complior device ®, ALAM medical; SphygmoCor ® device, AtCor medical).
Standardization of methods for comparison of devices
Because of the expansion of the field for non-invasive assessment of vascular function, devices are being constructed with varying pulse sensing techniques and signal processing algorithms. For proper and useful comparison of devices, there is a need for standardization of procedures and protocols. Such activities generally come from learned societies in the form of “Guidelines”. For comparison of PWV devices, the Society for Artery Research has published specific guidelines for device validation (294). There are tables for sample size (90 subjects selected with a minimum of 83 for data analysis), age range (at last 25 in age range <30, 30–60, >60 years) and exclusion criteria (e.g. BMI > 30 kg/m2, absence of sinus rhythm, significant arterial stenosis). There is also a specific description of the order of measurements between the devices so as to avoid the possibility of systematic errors. The results of device/method validation studies should be presented using the method of Bland and Altman (295), where the difference between the values obtained with the two devices is plotted against the mean value of both devices. The plot then shows the mean of, and the difference between, the two methods/devices and includes +/− 2 standard deviations as boundaries. Excellent, acceptable and poor accuracy may be defined as shown in Table 2 (294). Moreover, any systematic bias with respect to one method will be evident from the plot. Special consideration should be given to the issue of TD estimation, as different estimations between the devices will result in systematic over- or underestimation of cfPWV.
This protocol was recently used for the first time to validate a cuff based device (SphygmoCor XCEL ®) for detection of carotid femoral pulse transit time, with the aim to provide similar cfPWV values as those obtained with a femoral tonometer (25). When the cuff measurement of pulse transit time was corrected for the distance between the femoral site and the position of the cuff on the upper thigh, both devices gave similar cfPWV (R2 = 0.9) with mean difference of 0.02 m/s and SD of 0.61 m/s.
The problem of non-invasive estimation of travel distance for cfPWV measurement
In the measurement of cfPWV, the major source of inaccuracy lies in the determination of the TD of the pressure or flow waves (296). First, measurements on body surface may not reliably represent the true length of the aortic and arterial segments, especially with obesity and when the arteries become more tortuous with age (297). Second, by definition, cfPWV encompasses not only the aorta, but also segments of the carotid artery and of the iliac and femoral arteries, which differ with respect to their elastic properties (and their local PWVs) from the aorta, even more so during aging. Moreover, the proximal part of the aorta (the most elastic one), which undergoes marked changes with aging (297), is not covered. Finally, by definition, cfPWV encompasses the travel of the pulse wave up to the carotid artery and down the thoracic aorta at the same time. This is, thus, not a simple unidrectional path length (137), thereby rendering all determinations of the “real“ travelled path length somewhat elusive. Even sophisticated MRI-based distance measurements are valid only based on the assumption that the velocities in the carotid artery and in the thoracic aorta are the same, which actually may not be the case. In animals, PWV in the carotid artery can be 2–3 m/sec higher than in the aortic arch (298), and in humans the differences between aortic and carotid stiffness are higher in patients with hypertension and diabetes (299). Whether these differences can affect the actual cfPWV by 2% or up to 10% has been recently discussed (300). However, some standardization is obviously necessary, and comparisons of cfPWV with invasive PWV and MRI-PWV have been made. In 135 patients undergoing invasive coronary angiography, the “subtraction method“ (suprasternal notch (SSN)-femoral artery minus SSN-carotid artery) resulted in the smallest differences (0.2 m/sec) between invasive aoPWV and non-invasive cfPWW (14), whereas the direct distance method overestimated aoPWV by 2.9 m/sec. When the same TT (carotid-femoral TT derived from tonometry) was used, and TD was measured with MRI (aortic arch to femoral recording site minus carotid length from origin to recording site; again assuming equal velocities in carotid artery and aortic arch), the surface measurement closest to MRI TD estimate was carotid-femoral minus SSN-carotid (297). In another study, using MRI as reference for TD measurement (ascending aorta-femoral artery minus ascending aorta-carotid artery), the best estimate, as measured on body surface, was carotid-femoral distance multiplied by 0.8 (301). In all 3 studies, the direct carotid-femoral measurement led to a substantial overestimation of aortic PWV. Although conversion factors between the different cfPWV values obtained with different methods to assess TD are available from collaborative projects (139), this panel recommends to use either the “subtraction method“ (suprasternal notch – femoral recording site minus suprasternal notch – carotid recording site) or the “80 % method“ (80% of the measured direct distance between carotid and femoral recording site) to estimate TD for cfPWV. Additionally, the use of calipers may improve distance measurements particularly in overweight or obese subjects (302;303).
A comparison between different methods and devices, accuracy, repeatability, and reproducibility are summarized in the Online Supplement Section 6. In addition, a section summarizing clinical validation, i.e. which devices and techniques have been used in longitudinal clinical studies, again with a table, is summarized in the Online Supplement Section 6. Finally, a more-detailed discussion of devices that provide an estimate of PWV from waveform analysis or local arterial stiffness are also in the Online Supplement Section 6.
Validation of devices to measure brachial-ankle Pulse Wave Velocity (baPWV)
Repeatability and reproducibility can be investigated as usual, and such studies have been performed successfully (36;44). TD for baPWV can obviously only be estimated, as there is of course no direct unidirectional propagation of pressure or flow from brachial to ankle. The formula used in the systems is based on anthropometric data from Asian individuals, which may differ from Western populations. Although the travelled path with baPWV clearly differs from pure aortic (invasive) PWV and also from cfPWV through the inclusion of longer segments of muscular arteries, comparisons with invasive PWV (36) and cfPWV (304) have been made, showing a high degree of correlation. For non-invasive validation studies, systems that have been shown to predict cardiovascular outcomes should be used such as the VP1000 ® (Omron Healthcare, Japan) and the Vasera (Fukuda Denshi, Japan) (see Section 2).
Validation of devices providing estimates of PWV from single-point measurements
There is some interest in techniques estimating aortic PWV from brachial cuff-based waveforms analysis (and clinical characteristics), which would simplify the procedure. In addition to reproducibility, such devices should undergo invasive validation, when claiming to estimate aortic PWV, and/or non-invasive validation against gold-standard devices measuring cfPWV. So far, invasive validation has been performed successfully for the Arteriograph © (Arteriomed, Hungary) (305) and the Mobil-O-Graph © (I.E.M., Germany) (40). Clinical validation, i.e. the prediction of cardiovascular events, is pending for the Arteriograph. One small study in patients with chronic kidney disease, NKF stages 2–4, has already shown the independent prognostic value of an estimated aortic PWV (measured with the Mobil-O-Graph device) with respect to mortality (306).
Section 7. Factors confounding arterial stiffness measures and practical interpretation of values
Recommendations
7.1 Mean arterial pressure (MAP) and heart rate should be recorded at the time of an arterial stiffness measurement and taken into consideration when analyzing PWV data as potential confounders (Class I, Level of Evidence B) (15;16).
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7.2 The following are recommendations to enhance uniformity in arterial stiffness investigations:
i. Clearly state the sites of measurement in the Methods section, for example “carotid-femoral” (Class I, Level of Evidence C);
ii. It is reasonable to report how the distance measurement was performed in the Methods section (Class IIa, Level of Evidence C);
iii. It is reasonable to use calipers to obtain surface measurements to calculate distance for PWV (Class IIa, Level of Evidence C);
iv. We recommend that you perform arterial stiffness measurements, in duplicate, supine after a minimum of 10 minutes of rest, controlling the environmental noise and temperature as much as possible; you should repeat the arterial stiffness measurement a third time if the difference in the two measurements is more than 0.5 m/sec using the median value (Class I, Level of Evidence C);
v. You should ensure that operators performing arterial stiffness measurements are familiar with the equipment, have been trained in the techniques, and have demonstrated consistently reproducible results (Class I, Level of Evidence C).
A number of physiological and methodological factors can influence and confound arterial stiffness indices. These factors require due consideration in order to minimize their impact, allow high quality data to be obtained, and allow correct interpretation of the data.
Physiological confounders
The most significant physiological variable affecting arterial stiffness is the vessel distending pressure (mean arterial pressure; MAP) (15;18;307;308). In contrast, pulse pressure provides an indirect index of large artery stiffness because it depends on large artery compliance, together with stroke volume and the influence of reflected pressure waves. As MAP increases, vessels stiffen, but in a non-linear manner. Therefore, the measured value of stiffness will depend on, or be confounded by, the mean arterial pressure, which should be taken into consideration. This is particularly relevant when comparing populations with different blood pressures or when investigating the effects of anti-hypertensive agents.
The relationship between heart rate and arterial stiffness is less well defined, with acute studies showing positive associations (16;309;310), no association (311;312), or even inverse associations (313) between increased heart rate and various measures of arterial stiffness, including pulse wave velocity. These disparate results reflect that at least some of the studies may have been confounded by concomitant changes in MAP. Nevertheless, a recent study (314) demonstrated that although heart rate exerts a minimal influence on pulse wave velocity in the lower range of mean pressure values, an increase in heart rate results in a modest but significant increase in pulse wave velocity at higher MAP values. Since blood pressure and heart rate vary considerably both within and between individuals, both should be taken into consideration when undertaking measurements of arterial stiffness.
To minimize such confounding effects arterial stiffness should be assessed in a quiet, temperature-controlled environment. Participants should also refrain from alcohol, vasoactive medications, and vigorous physical activity ideally for 12 hours, and large meals, caffeine-containing food and drinks and smoking for at least 2–4 hours prior to the measurements. It is important that participants are allowed to rest, in the supine position for at least 10 minutes to ensure hemodynamic stability. For menstruating women attention should be paid to studying these subjects at a similar menstrual cycle phase.
Methodological confounders
Although cfPWV is recognized as the gold-standard for the non-invasive assessment of arterial stiffness (45), often arterial stiffness is measured in alternative (or additional) vascular beds. For example, several noninvasive commercial devices assess baPWV. Compared to the carotid-femoral vascular bed, the brachial-ankle vascular bed encompasses additional arterial territories with different characteristics, different determinants of stiffness, and different influences of atherosclerosis. Conversely, invasive assessments of arterial stiffness and MRI-guided assessments of arterial stiffness often measure PWV across much shorter distances within the aorta. It is important to recognize that indices are not necessarily interchangeable, either physiologically or prognostically, and it is important to clearly state the methodology used to assess PWV.
Even within a vascular bed, PWV may vary depending on the specific device that is used to measure PWV. For example, Millasseau et al (315) assessed cfPWV with 2 commercially available devices in the same individuals. They found that the 2 devices yielded different values of PWV within the same individual. Importantly, the difference was attributable to the algorithm used by each device to derive the time of travel (foot-to-foot method with the SphygmoCor © system vs. maximum slope method with the Complior © system), such that the same waveforms analyzed by the 2 devices could result in differences in PWV values of 5–15%.
Perhaps the most important methodological confounder of pulse wave velocity measurements is calculation of the wave travel distance (see Section 6). cfPWV is calculated as the distance travelled by the pressure wave divided by the time delay between the arrival of the pulse wave at the carotid and femoral sites (wave transit time). For measurement techniques other than MRI, the travel distance is typically estimated from surface measurements between the recording sites. These measurements should be as accurate as possible, since small errors in distance measurement may translate into much larger errors in the calculated pulse wave velocity - up to 30% in one study (316), and the measurement method and vascular territory should be clearly stated in the manuscript.
A tape measure is generally used, although calipers better minimize the impact of body contours and, therefore, are recommended. Different approaches are used to calculate wave travel distance, although the most common methods are the direct distance between the carotid and femoral sites (direct method) and the distance between the suprasternal notch and carotid site subtracted from the distance between the suprasternal notch and the femoral site (subtracted method), which better corresponds to the true anatomical distance assessed by MRI (301). Weber et al also found that the subtraction method was more closely related to true distance and that carotid-femoral PWV determined with the device and the subtracted distance corresponded best with invasive assessment of PWV (14). Although a recent expert consensus document advised distance should be calculated by multiplying the direct distance by 0.8 and conversion algorithms between the two methods have been developed (317), these are likely to introduce further error. Therefore, the method of distance calculation should be clearly stated, and the subtracted distance is more anatomically true (see section 6 recommendation). How the application of different methodologies will relate to differences in risk prediction remains unclear.
Practical consideration in making arterial stiffness measurements
Whenever tonometry or ultrasound systems are used for sequential recording of pressure or flow waves, using ECG gating, care has to be taken that cardiac rhythm is stable. In the presence of arrhythmias, measurements may be unreliable due to different intervals from ECG’s R-wave to the foot of the travelling wave.
In addition to physiological and other confounders of arterial stiffness measurements, there are a number of limitations associated with assessing arterial stiffness. Some of the techniques are highly operator-dependent and thus adequate training for the individuals making the recordings must be provided to ensure that high quality data are obtained. Therefore, a period of familiarization with the measurement techniques is suggested, after which the trainee should obtain high quality recordings in a minimum of 20 individuals to determine competency. In addition, the equipment required for these measurements is often expensive and lacking in portability, limiting the use of some techniques for measuring arterial stiffness to specialist research settings. This is especially the case for MRI- and ultrasound-based approaches, although a number of portable ultrasound systems are now available.
Section 8: Future Needs in Arterial Stiffness Study
Understanding how aging, stiffness and BP interact over time is a complex conundrum. “Aging”-associated arterial changes, and those associated with hypertension (and early atherosclerosis and diabetes), are fundamentally intertwined at the cellular and molecular levels. In humans, other well-known risk factors (e.g., excess food intake, altered dietary lipid and metabolism, smoking, and lack of exercise) likely interact with this arterial substrate that has been altered during aging, and that renders the aging artery a “fertile soil” that facilitates the initiation and progression of these arterial diseases. Some lifestyle and pharmacologic interventions have already proved to be effective in preventing or ameliorating hypertension associated with aging. Although a number of small studies have suggested that various life style interventions may produce BP independent decreases in carotid-femoral PWV, to date the best evidence available in terms of therapeutic intervention suggests that ACE inhibition may produce decreases in arterial stiffness beyond a blood pressure lowering effect (318;319). Much larger meta-analyses of individual patient data will be required in the future to be sure that decreases in aortic PWV following therapy are truly blood pressure independent. The cellular/molecular pro-inflammatory mechanisms driven by Ang II and other growth factors that underlie arterial aging are novel putative candidates to be targeted by interventions aimed at attenuating arterial aging, and thus possibly attenuating the major risk factor for hypertension and atherosclerosis (320).
Future needs in investigating the import of arterial stiffness should include addressing questions such as:
Do age changes, per se, within the arterial wall drive the age-associated increase in arterial stiffness, or does the increase in arterial stiffness with advancing age result from the age-associated increase in systolic BP?
What is the natural history of PWV and BP, vis-à-vis the rate at which PWV and BP increase with age?
Will prevention, or reduction, of aortic stiffening provide substantial health benefits?
What are the targets for intervention in a focused attempt to alter the nature of the arterial wall?
Is it possible, and is it safe, to unstiffen the aorta independent of blood pressure reduction?
Can the similarities in aging and stiffening of the arterial wall in animal models be used to guide human intervention trials; will Industry or Peer Review organizations consider these processes as potentially tractable and fund investigations into intervention trials? How would such trials differentiate the impact of a de-stiffening approach from a reduction in blood pressure?
Are there non-drug interventions that are likely to benefit arterial stiffening processes? At what age should such interventions be introduced?
Many of the above will be facilitated by the development of cuff based systems which will allow for the measurement of hemodynamic parameters such as cfPWV, central blood pressure and AIx with as much ease and operator independence as oscillometric sphygmomanometry. Such systems have already been validated, and, in addition, have the facility for 24hr ambulatory assessment of central blood pressure (such as the Mobil-O-Graph © covered in Section 2). A logical progression would be to measure cfPWV using non cuff based systems and such systems are already in development (321).
The establishment of international reference norms for PWV across age and blood pressure strata (139), increasing recognition of the importance of central arterial stiffness as a consequence of aging and co-morbidities (22;105), potential improvements in understanding study outcome mechanisms when these measurements are incorporated (59;196), recognition of the limitations of these measurements coupled with a spirit of cooperation between Device Manufacturers, the Pharmaceutical Industry, Regulatory Sponsors, Payers, Investigators, Practitioners and Patients are necessary foundational elements in moving this process forward.
In addition there are several gaps in the understanding of arterial stiffness in children. These include:
Lack of validation of measurement devices in children
Lack of sufficient normative data by age/body size/pubertal status, gender, race
Lack of longitudinal data in healthy children, and children with risk factors (diabetes, hypertension)
Linking arterial stiffness measurements to established pediatric intermediate target organ endpoints
As this summary statement was nearing final draft stage a large patient-level (n=17,635) meta-analysis of arterial stiffness was published (3). This lends more support to the growing interest in arterial stiffness.
Overall Summary
Measuring arterial stiffness has been established clinically through longitudinal studies where it has independently predicted death, as well as standard cardiovascular endpoints. A number of devices and approaches have been developed to assess this parameter providing both challenges and opportunities for the advancement of this aspect of the science of hemodynamics. Wider appreciation of the role of arterial stiffness beyond blood pressure levels in clinical medicine and clinical research is an ongoing journey, and its indication for use in the clinic requires further study. The authors hope this Summary Statement represents a forward step in this journey.
Supplementary Material
Abbreviations used in this Statement (unit of measure, when applicable)
- ACE
Angiotensin converting enzyme
- ARB
Angiotensin Receptor Blocker
- AIx
Augmentation Index – a ratio expressing the relationship of forward and backward traveling waves in the central aorta (unit-less, or sometimes expressed as %)
- Ang
Angiotensin
- BP
Blood Pressure
- baPWV
Brachial-Ankle Pulse Wave Velocity (a measure of arterials stiffness; meters/second)
- cfPWV
Carotid-Femoral Pulse Wave Velocity (a standard measure of arterial stiffness; meters/second)
- CKD
Chronic Kidney Disease
- CV
Cardiovascular
- CVD
Cardiovascular Disease
- DM
Diabetes Mellitus
- eGFR
Estimated Glomerular Filtration Rate (typically from the MDRD equation; mL/min/1.73m2)
- ECG
Electrocardiogram
- Ea
Arterial Elastance (a measure that relates end systolic pressure to LV stroke volume)
- EECP
Enhanced External Counter-Pulsation
- Einc
Elastic Modulus (dynes/cm2)
- ESC
European Society of Cardiology
- ESH
European Society of Hypertension
- ESRD
End Stage Renal Disease
- Ew
Wasted left ventricular workload energy
- fdPWV
Femoral-Dorsalis Pedis Pulse Wave Velocity (a measure of arterials stiffness; meters/second)
- GFR
Glomerular Filtration Rate (a measure of kidney function; mL/minute)
- HF
Heart Failure
- IDI
Integrated Discrimination Improvement (statistics procedure)
- LDL
Low Density Lipoprotein
- LOE
Level of Evidence
- LV
Left Ventricle
- LVH
Left Ventricle Hypertrophy
- LVOT
Left Ventricle Outflow Tract
- MAP
Mean Arterial Pressure (mmHg)
- MRI
Magnetic Resonance Imaging
- NIR
Net Reclassification Index (ratio)
- PC-MRI
Phase Contrast Magnetic Resonance Imaging
- PP
Pulse Pressure (systolic minus diastolic pressure; mmHg)
- PWA
Pulse Wave Analysis – use of an arterial waveform to interrogate vascular function (units vary)
- PWV
Pulse Wave Velocity – the standard measure of arterial stiffness (meters/second)
- RM
Reflected wave Magnitude (mmHg)
- SBP
Systolic Blood Pressure
- SD
Standard Deviation
- SVR
Systemic Vascular Resistance (dynes-second/cm5)
- TD
Travel Distance (millimeters or meters)
- TT
Transit Time (milliseconds, or seconds)
- VENC
Velocity-Encoding Sensitivity (an MRI technique to measure flow velocity)
- Zc
Characteristic aortic impedance (also called Zo; measures aortic pressure-flow relationship; ((mmHg*sec)/mL)
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