Eur J Cardio-thorac

Surg (1992) 6: 609-617

swrgery

0 Springer-Verlag 1992

Quantitation of the turbulent stress distribution downstream of normal, diseased and artificial aortic valves in humans H. Nygaard ‘7‘, P. K. Paulsen’, J. M. Hasenkam ‘*3, 0. Kromann-Hansen ‘, E. M. Pedersen ‘*3, and P. E. Rovsing ‘*’ ’ Department of Thoracic and Cardiovascular Surgery, Skejby Sygehus, Aarhus University Hospital, Brendstrupgardsvej, DK-8200 Aarhus N, Denmark * The Engineering College, Aarhus Teknikum, Dalgas Avenue 2, DK-8000 Aarhus C, Denmark 3 Institute of Experimental Clinical Research, Skejby Sygehus, Aarhus University Hospital, Brendstrupgirdsveij, DK-8200 Aarhus N, Denmark

Abstract. Damage to blood corpuscles seems to be related to the magnitude and exposure time of the turbulent shear stresses (TSS). According to in vitro studies the critical TSS level for lethal erythrocyte and thrombocyte damage is 150-400 N/m*, for exposure times within physiological ranges. To study the distribution of TSS in the human ascending aorta, a hot-film anemometer needle probe was used to register blood velocities at 41 evenly distributed measuring points in the cross-sectional area 5-6 cm downstream of the aortic ammlus. Measurements were made in the ascending aorta after normal aortic valves (prior to coronary bypass surgery), after stenotic aortic valves, and after implantation of either St. Jude Medical or Starr Edwards Silastic Ball valves. Three-dimensional visualization of velocity profdes were performed and Reynolds normal stresses (RNS) were calculated within 50-ms overlapping time windows in systole. By coordinating the mean RNS for each time window and for all 41 measuring points, 2-dimensional color-coded mapping of the RNS distribution was made. Based on the velocity profiles and the RNS distribution a relative blood damage index (RBDI) was calculated to incorporate the magnitude and exposure thne for RNS in the entire cross-sectional area into one parameter. Turbulent shear stresses were estimated by using a previously determined correlation equation between RNS and TSS. After normal aortic valves, RNS was below 4 N/m’. This is an order of magnitude less than RNS after the stenotic valves which was up to 38 N/m’. After artifical aortic valves, RNS up to 120 N/m’ were found. The highest RNS values were generally found in areas with high or rapidly changing velocity gradients. The estimated RBDI emphasized the difference between normal, stenotic and artificial heart valves, revealing an order of magnitude between the three categories and showing no significant difference between the two investigated prosthetic valve types. By measuring blood velocity and turbulence downstream of aortic valves, the degree of blood cell damage caused by the valve can be estimated. The levels of RNS and RBDI found for the two artificial valves are sufficient to cause sublethal damage to blood cells. [Eur J Cardio-thorac Surg (1992) 6:609-6171 Key words: Humans - Native aortic valves - Prosthetic aortic valves - Turbulent stresses - Hemodynamics anemometry

Introduction Diseased heart valves have now been replaced by prostheses for more than 3 decades, generally leading to improved survival and enhanced quality of life. However, some clinically important late valve-related complications still exist. These include valve thrombosis, thromboembolic events, hemolysis, tissue overgrowth, calcification, prosthetic valve endocarditis and mechanical failure. These complications account for approximately half of the late deaths [37]. Most of these problems seem related to the fluid dynamic characteristics of the valve, which Presented at the Poster Session of the 5th Annual Meeting of the European Association for Cardio-thoracic Surgery, London, UK, September 23-25, 1991 Correspondence to: Prof. Hans Nygaard

- Hot-film

include: high turbulent stresses, regions of flow reversal or flow separation, regions of relative stasis, high local velocity gradients and vibrations due to eddy formation. Turbulence was one of the first features of flow to gain attention as a potential factor associated with a variety of pathophysiological changes such as endothelial damage [5,21], platelet damage [12,15,38,40,46], and hemolysis [l, 16, 20, 35, 421. These studies revealed a relationship between the degree of endothelial or blood cell damage and the magnitude and exposure time of wall or bulk shear stresses. The evaluation of turbulence in the cardiovascular system is complex, due to many variables not commonly encountered in engineering applications including: pulsatile flow, non-Newtonian fluid, compliant vessels, varying contractility, complex entrance regions, tapering of vessels, curvatures and branches of vessels.

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All prosthetic heart valves are stenotic to some extent. In the normal human aortic valve almost the entire area of the annulus is available for blood flow. When an artificial heart valve is inserted, the area available for blood flow is reduced by the area taken up by the sewing ring and any mechanical or biological devices within the sewing ring. This evidently disturbs the flow in the near vicinity of the valve, producing turbulence and recirculating areas as well as regions of stasis. Numerous in vitro studies have clearly demonstrated turbulence or highly disturbed flow downstream of different types of artificial heart valves [2, 3,7,9, 10,491. In animal studies turbulent blood flow has been observed in the ascending aorta of dogs [4, 22, 36,481, and quantitative studies of turbulent stresses downstream of biological and mechanical aortic valve prostheses have been performed in pigs [ll]. Beyond clinical follow-up studies [14, 17- 19, 34, 401, human experimental studies have been limited. The quantitative data available on turbulence in the ascending aorta, therefore, has been obtained in vitro and in animal studies. As however, results from these studies do not always correlate with clinical observations, it is pertinent to expand the data to comprise quantitative human studies on turbulence as well. Accordingly, the main objective of this study was to gain a more quantitative understanding of the temporal and spatial development of turbulent stresses during the systole in patients with normal, stenotic and artificial aortic valves in relation to the damaging effect such stresses may have upon red blood cells and platelets. This was achieved by three-dimensional visualizations of the velocity profiles and by two-dimensional color-coded mapping of the turbulent stress distribution, both types of data shown as a function of time during the systolic ejection phase. From these results, a relative blood damage index (RBDI) was calculated to combine the magnitude and exposure time of the turbulent stresses in the entire cross-sectional area.

Theoretical considerations Hinze [13] defined turbulence as “motion in an irregular condition of flow in which the various quantities show a random variation with time and space coordinates so that statistically significant average values can be discerned”. According to this strict definition, the presence of turbulence can be proven only by simultaneous registration of random velocity fluctuations in three dimensions. In practice, however, the process has been so well characterized in many fluid dynamic studies that random fluctuations in one plane can be accepted as evidence of turbulence [22, 231. The fluid stresses associated with turbulence, the socalled Reynolds stresses, can be interpreted as stresses that acts on an element of fluid. The turbulent stress acting normally on an element of the fluid is called Reynolds normal stress (RNS), and is defined as: ,z RNS= QU,, where u’, is the turbulent velocity component in the axial direction and Q the fluid density.

The turbulent shear stress (TSS) or Reynolds shear stress is defined as: TSS = &u; and can be interpreted as the stresses acting tangentially on a surface of an element of the fluid. ufi and u’~ are orthogonal turbulent velocity components recorded simultaneously. Turbulent shear stress can be measured directly in vitro using a two-component Laser Doppler anemometer (LDA) system [7,44,45], but is so far impossible to measure in vivo, due to lack of feasible methods. Reynolds normal stress, on the other hand, can be calculated from a single velocity component which can be measured in vivo using hot-film anemometry (HFA) [l 1,411. Since damage to blood corpuscles seems associated with the magnitude and exposure time of TSS, some investigators estimated TSS from measurements of a single axial velocity component [33]. These estimates were based on theoretical considerations by Tennekes and Lumley [43]. They suggested a correlation factor between RNS and TSS of 0.4 for steady two-dimensional homogeneous turbulent shear flow. Hanle et al. [9] used the same factor, but they measured the radial as well as the axial velocity components separately with a one-component LDA system. Nygaard et al. [24] performed a correlation analysis between RNS and TSS in a pulsatile flow model downstream of prosthetic aortic valves. It was concluded that the maximum turbulent shear stresses could be estimated as: TSS x0.6 RNS, for HFA measurement of RNS. Many research groups have studied the relation between critical shear stress levels and their exposure times in causing hemolysis and platelet lysis. The results of these investigations were reviewed by Giersiepen [6], according to whom the critical shear stress level for lethal erythrocyte and thrombocyte damage is 150-400 Nm-‘, for exposure times within physiological ranges. In order to interpret these data, Giersiepen et al. [8] developed a mathematical model for determination of shear stress-induced blood cell damage, based on in vitro studies of the mechanisms of damage and the effects of shear stresses on blood corpuscles [46,47]. According to this model the relative lactic dehydrogenase (LDH)-release by platelets and the relative hemoglobin release by damaged red blood cells can be expressed as a function of TSS and its exposure time in a certain measuring area. These relations were integrated spatially along one vessel diameter and temporally through the entire systole, and then added to calculate a relative blood damage index (RBDI). The model was evaluated by calculating RBDI from in vitro velocity measurements downstream of 25 different types of mechanical and bioprosthetic heart valve prostheses and correlating it with clinical long-term studies of Horstkotte et al. [14]. Comparison of the experimentally determined RBDI with clinical evidence of hemolysis generally revealed a good correlation. In the present study, the calculation procedure of RBDI was based on a slightly modified version developed by Nygaard et al. [25] to include the entire cross-sectional area of the vessel. Thus RBDI takes into account both the magnitude of TSS, its exposure time and the area

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in which this load has been effective. The turbulent shear stress was estimated by the previously determined correlation equation between RNS and TSS.

Material and methods A l-mm L-shaped HFA-probe operated with a 55 M 01 Main Unit, 55 M 10 Standard Bridge, and 55 M 25 Linearizer (all Dantec equipment) was used for point blood velocity measurements. The tip of the probe was positioned intraluminally 5-6 cm downstream of the aortic valve, through a purse string suture in the anterior vessel wall as indicated in Fig. 1. Measurements were performed at 41 almost evenly distributed measuring points in the cross-sectional area using a specially designed probe positioning device [28]. The principle of measurement is that a thin metal film mounted on the tip of the probe is heated up 5°C above the surrounding blood temperature. The passing blood tends to cool off the film by convection and the power required to maintain the film temperature is an expression of fluid velocity at the measuring point. Since the relation between velocity and output voltage from the system is unlinear, linearization is necessary. Calibration and linearization procedures have been described previously [26,27]. The system was equipped with a specially constructed patient safety unit [29], to prevent electric shock and thermal damage to the blood. The ECG and linearized HFA velocity signals were recorded on an instrumentation recorder (Hewlett Packard, type 3960) for later data analysis. Further details about the operative technique and anesthesia were described by Paulsen et al. [31]. Data from that investigation was reanalyzed for the present study and supplemented with measurements on patients with aortic valve stenosis, using the same procedure. This study comprised 17 patients (8 females and 9 males). The mean age was 56 years (range 42-72 years). Measurements were performed in 8 patients with normal aortic valves prior to coronary bypass surgery, in 2 patients with stenotic aortic valves prior to cardiopulmonary bypass, in 4 patients after insertion of a St. Jude Medical (SJM) aortic valve, and in 3 patients after insertion of a Starr-Edwards Silastic Ball (SSB) aortic valve. The measurements were conducted in a stable hemodynamic environment, and lasted 15-20 min. The study was approved by the Danish Local Ethical Committee.

Signal processing In each of the 41 measuring points 5- 15 heart cycles of the instantaneous velocity signal were recorded. The systolic part of the veloc-

I

Hot-film -I ;;tbtmeter

Positioning

Vessel

ity signals was averaged with reference to the ECG to calculate one mean systolic velocity curve for each measuring point ( in Fig. 2). A computerized 3-dimensional plot of the spatial distribution of blood velocity during systole was made by coordinating the mean blood velocity from all measuring points at specific times in the cardiac cycle. The procedure was explained in detail by Nygaard et al. [25]. The turbulent stresses were calculated within eight 50-ms overlapping time windows during systole, for each of the 41 measuring points. Due to the small number of heart cycles in the first measuring series, the turbulent velocity component u’(t) was extracted from the instantaneous velocity component u(t) by 20 Hz high-pass filtering (Fig. 2). A signal analyzer (Briiel & Kjsr, Dual Channel Signal Analyzer, type 2032) was used to calculate the mean axial turbulence energy density spectrum for each of the 41 measuring points and RNS was then calculated by multiplying the mean axial turbulence energy by blood density. By coordinating the RNS for each time window and for all measuring points, a computerized 2-dimensional color-coded mapping of the turbulent stress distribution was made as described in detail by Nygaard et al. [25]. Based on the temporal development of the velocity profiles and turbulent stress distribution, RBDI was estimated for all patients.

Results The measured and calculated key parameters for the four patient groups are listed in Table 1. The average of the 41 peak velocities (VP& was generally lower in patients with normal aortic valves than in patients with an artificial valve. This tendency is also reflected in the peak Reynolds numbers (Re,,,,). One of the patients with an aortic valve stenosis had peak velocities and peak Reynolds numbers of the same order of magnitude as the patients with normal valves, while the other patient with a stenotic valve had values more like those of the patients with artificial valves. For normal valves, maximum RNS was below 4 N/m’ and the spatial averaged mean systolic RNS was below 2 N/m’, which is an order of magnitude lower than in the patients with stenotic valves, where maximum and mean RNS were up to 30 N/m2 and 7 N/m2 respectively. The highest values of RNS were found in patients who had an artificial aortic valve implanted. The maximum and mean values of RNS for patients with an implanted

Device

Fig. 1. A diagram of the hot-film probe sup porting and positioning device (left and the distribution of the 41 measuring points in the cross-sectional area of the vessel and relations to sinuses of valsalvae (right). The measuring points in the vessel lumen could be reached by the tip of the hot-film probe by regulating depth and angle on the positioning device

612 time windows

Turbulent

Underlying

Table 1. Measured and calculated key parameters Patient no.

Nl -N8 AS1 AS2 Jl 52 53 54 Sl s2 s3

Measurement

situation

Before coronary bypass Before valve implantation Before valve implantation SJMZI SJMZI SJM23 SJM21 SSBIZ SSBII SSBIZ

Fig. 2. A schematic presentation of the signal processing. The instantaneous velocity u(t) was subjected to ensemble averaging, to calculate one mean velocity curve for each measuring point (u(t)). The turbulent velocity component u’(t) was extracted from the instantaneous velocity component by 20 Hz high-pass filtering, to calculate the turbuient stresses within SO-ms overlapping time windows (dotted lines) during systole

velocity

velocity

for the four patient groups V*car (cm/s)

Repelk

56-85 73 150 108 117 130 180 162 110 160

2 947-5 158 4 610 8 624 5 115 7 389 5 473 8 526 6 821 6 368 IO 947

$Z)

RN%,,, (N/m2)

RBDI,,, x 10-2