Arterial pressure waveform was continuously recorded at the left middle finger by Finometer system (Finometer MIDI, Finapres Medical Systems, Amsterdam, Netherlands) during WWI and stored on a computer using a data acquisition system (PowerLab, AD Instrument) at the 1000 Hz of sampling rate. Throughout the WWI, subject was asked to keep the left hand at the heart level: on the side-table out of the bathtub. SV, CO, and total peripheral resistance (TPR) were estimated with a non-invasive blood pressure (BP) measurement device incorporated Modelflow-based hemodynamics measurement software (Finometer Model-2, Finapres Medical Systems, Amsterdam, Netherlands). The validity of the Modelflow method to derive hemodynamic measurements has been established in a variety of conditions (Wesseling et al., 1993; Sugawara et al., 2003). HR, brachial BP, and PWV were measured with a vascular testing device equipped with an electrocardiogram, phonocardiogram, oscillometric extremities cuffs (form PWV/ABI; Colin Medical Technology, Komaki, Japan), and an applanation tonometry sensor unit (TU-100; Colin Medical Technology, Komaki, Japan), as previously described (Kosaki et al., 2015; Sugawara et al., 2016, 2019). Carotid and femoral arterial pressure waveforms were simultaneously recorded by two applanation tonometry sensors incorporating an array of 15 micro-piezoresistive transducers. Briefly, PWV (=arterial path length/pulse transit time) was obtained between the carotid and femoral regions (e.g., aorta) and between the femoral and ankle regions (e.g., leg) as an indirect index of local arterial stiffness. Arterial path lengths were assessed with a straight distance measurement over the surface of the body using a steel measure. The carotid-femoral body surface straight distance multiplied 0.8 was applied for aortic PWV calculation (Van Bortel et al., 2012).

Aortic pressure waveforms were synthesized from the carotid pressure waveforms with the generalized transfer function using SphygmoCor system (Model MM3, AtCor Medical, West Ryde, NSW, Australia). Measurements of aortic pulse wave analysis are presented in Figure 1. Carotid and aortic BP were calibrated with the brachial mean arterial pressure and diastolic BP (Armentano et al., 1995). Aortic augmentation pressure (AP) was defined as the difference between the first systolic peak (or shoulder) and the second systolic peak pressures (Kelly et al., 1989). Aortic augmentation index (AIx) was calculated as the percent ratio of aortic AP to aortic pulse pressure and standardized for an HR of 75 bpm. From the synthesized aortic pulse, areas under the systolic and diastolic parts of the curves [the time-tension index (TTI) and diastolic pressure-time index (DPTI)] were obtained. Subendocardial viability ratio (SEVR = DPTI/TTI100) was calculated as the ratio of the myocardial oxygen demand and the blood supply to the heart (Buckberg et al., 1972; Tanaka et al., 2016).

Schema of aortic pulse wave analysis. Augmentation pressure (AP) is the additional pressure added to the forward wave by the reflected wave. Augmentation index (AIx) is defined as the AP as a percentage of pulse pressure. The dicrotic notch (DN) represents the closure of the aortic valve and is used to calculate ejection duration. Aortic tension-time index (TTI) and diastolic pressure-time index (DPTI) reflect the myocardial oxygen demand and the blood supply to the heart, respectively. These are calculated as the areas under the curve of aortic pressure waveform in systole and diastole, respectively. The subendocardial viability ratio (SEVR) is the ratio of DPTI to TTI.

Femoral artery blood flow was measured with an ultrasound machine equipped with a high-resolution multi-frequency linear-array transducer (14 MHz) (CX50xMATRIX, Philips Ultrasound, Bothell, WA, United States), as we previously reported (Sugawara et al., 2007). Briefly, the longitudinal two-dimensional and Doppler ultrasound images were consecutively obtained below the inguinal ligament, 2–3 cm above the bifurcation into the profundus and superficial branches. For mean blood flow velocity measurements, a probe insonation angle was calibrated to 60°. The sample volume was adjusted to cover the width of the common femoral artery to encompass the entire lumen of the vessel, and the cursor was set at mid-vessel. Five-second B-mode movies and 5-s Doppler images were stored at least two of each as a DICOM file for later offline analysis. To eliminate the inter-investigator variability, the same person analyzed ultrasound images with computerized image-analysis software (ImageJ, NIH) in a blind manner. The arterial diameter was determined by a perpendicular measurement from the media/adventitia interface of the near-wall to the lumen/intima interface of the far wall of the vessel. A mean diameter was calculated based on the relative time periods of the systolic (1/3) and diastolic (2/3) BP phases and was used to represent the cross-sectional area. Three measurements of arterial lumen diameter were taken per frame and averaged. Blood flow was calculated as (mean blood velocity) × (circular area) × 6 × 104 (with the constant 6 × 104 being the conversion factor from m/s to L/min). Data reported are the time averages of ≥10 measurements for all variables. Femoral arterial shear stress (SS) was calculated using the following equation: 8 × blood viscosity (assumed to be 0.035 dyn × s/cm2) × mean blood velocity/baseline diameter (at end-diastole) (Dhindsa et al., 2008).

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