Validation and applications of mitral prosthetic valvular areas calculated by Doppler echocardiography

VALVULAR

Htikl

Validation

DISEASE

and Applications of Mitral Valvular Areas Calculated by Doppler Echocardiography

Prosthetic

Jean G. Dumesnil, MD, George N. Honos, MD, Michel Lemieux, MD, and Jocelyn Beauchemin, RT

Doppkrechocaidiographyisusedinthenoninvasive evaluation of mRral valve prostheses using parameters hsretofore validated primarily for titive valves. Accordtigly, this study was designed to examine the +idRy and relative usef&sss of vaive gradisnt and area me=-=tsinagroup of29patients(l7women,Smen,meanage92* 8 years), 19 f 4 months after implantation oi differentsixes(29to31mm)ofagiventypeofbio~ pros&&. Areas obtained with both the continuity equation, udg stroke volume B in the left ventrkular otitftow tract, and the preuuehatftime method .are compared to known pros&&k areas derived from an in vRro hydraulii mklel. Areas cakuiated by the centinuity equaticarvelateweUwRhbvRroareas(r=0.92,standarderroroftheyestimate=O.lem2, p 02), and are above ths range of predicted vaiuekin69%ofcases. correlatsonrarealsofound between continuity aquetson areas and the peak and mean valvubr gradisnts (r = 0.59, p
appiied in this study, is a valid methed for assessing mRral valve bioprorthaie areas provided that there is no signiint aortic or mRral regurgitation. The prauve half-time method is not valid in .this situstion and grossiy 0verestiWtes the area in many &ases. The projected relations between in&Xadprorthstie~cUsnd pressum gradients may behetpfulfor~sselectiitoavoida patieni--8 mismatch. (AmJCardiol1990;65:1443-1448) From the Quebec Heart Institute, Quebec, Canada. Manuscript received October 26,1989; revised manuscript received February 5,1990, and accepted February 7. Address for reprints: Jean G. Dumesnil, MD, Quebec Heart Institute, 2725, Chemin Samte-Foy, Sainte-Foy, Quebec, Canada, GlV 4G5.

oppler echocardiographjl is currently used for the noninvasive evaluation of mitral valve prostheses. The parameters most often measured with thii technique are the peak and mean valvular gradients, as well as the effective orifice area.1-6 Validation of the latter has been derived mainly from comparison with hemodynamic data in patients with native valve disease and there has been little validation in patients with mitral valve prostheses. Moreover, recent studies have shown that there are limitations to the pressure half-time method, which is currently the most commonly used method to calculate mitral valve orifice area.7-9 Our goals were to assess the validity of Doppler echocardiographic measurements, particularly effective orifice area calculations, in patients with mitral valve prostheses, and to determine the implications of such information with regard to prosthetic performance and implantation criteria. For this purpose, effective mitral orifice areas were calculated with Doppler echocardiography in a group of patients with a specific type of bioprosthesis of varying sizes, using both the continuity equation and pressure half-time methods. These measurements were then compared to known prosthetic effective orifice area values derived from in vitro hydraulic model flow measurements provided by the manufacturer.

D

MmHODS

The study population consisted of 26 patients (17 women and 9 men) with a mean age of 62 f 8 years (range 44 to 75) who underwent mitral valve replacement at the Quebec Heart Institute between May 21, 1986, and September 23, 1987. Preoperative diagnoses included mitral stenosis (18 patients), mitral insufficiency (5 patients) or mixed mitral valve disease (3 patients). Two patients had severe coronary artery disease and underwent concurrent coronary artery bypass graft surgery. All inserted prostheses were Medtronic IntactTM bioprostheses model 705, and the sizes used were 25 mm (7 patients), 27 mm (11 patients), 29 mm (4 patients) and 31 mm (4 patients). Doppler echocardiographic examination was performed a mean of 19 f 4 months after surgery. All studies were performed by the same technician on a Hewlett-Packard Son&-500 ultrasound system and subsequently independently reviewed by the same investigator. Four patients had to be excluded because of technically inadequate studies. There was complete agreement THE AMERICAN

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TABLE

I In Vitro

and In Vivo Measurements

in Intact

In Vivo Measurements

Prosthesis Size (mm) 25 27 29 31

Mitral

by Doppler

Bioprostheses

of Various

Sizes

I

Echocardiography Regurgitation (No. of patients Incidence)

In Vitro EOA (cm*) 1.34 (1.19-1.48) 1.51 (1.31-1.71) 1.65 (1.43-1.86) 1.86 (1.66-2.05)

No. of Pts 7 11 4 4

EOA by CE (cm*)

EOA by PHT (cm*)

Peak Gradient (mm W

Mean Gradient (mm Hz3

1.36 f 0.08 (1.20-1.46) 1.53 f 0.09 (1.41-1.74) 1.63 f 0.08 (1.49-1.69) 1.81 ho.15 (1.54-1.91)

1.78 f 0.32 (1.26-2.25) 2.02 f 0.29 (1.54-2.48) 2.28 f 0.62 (1 A62.93) 1.87 f 0.23 (1.51-2.12)

16.2 f 3.6 (10.4-21.0) 11.0*3.1 (6.4-18.0) 8.6 f 1.5 (7.3-10.9) 9.0 f 3.2 (5.4-13.5)

7.8 f 2.4 (4.2-12.7) 5.4* 1.5 (2.4-7.1) 3.6 f 0.4 (3.1-4.0) 3.6f 1.1 (2.4-5.0)

Mitral l(l+),

[Severity],

Aortic 14%

2(1+).

29%

0

ql+),

30%

0

2(1+).

50%

0

0

Values are mean f standard deviation with range of values in parentheses. CE = continuity equation; EOA = effectw oriice area; PHT = pressure half-time.

in measurements between technician and investigator in 19 cases, whereas measurements were redone with a frame-grabber off the videotape in 7 patients. Measurements obtained in patients for the purpose of this study consisted of peak and mean mitral valvular gradients, stroke volume, cardiac output and prosthetic valve area by both the continuity equation and pressure half-time methods. Peak and mean gradients were obtained with pulsed- and continuous-wave Doppler recordings of the mitral jet velocity from the apical window and application of the modified Bernoulli equation as previously described and validated. lo-r2 Using color-flow Doppler as a guide, we took great care to ensure optimal alignment between the Doppler cursor and the mitral flow jet; an angle <20° was considered a minimal requirement and the maximal velocities recorded were used to determine the pressure gradients. Stroke volume was calculated from the left ventricular outflow tract area, as estimated from the corresponding internal diameter measured in the left parasternal long-axis view in systole, multiplied by the velocity-time integral of the left ventricular outflow tract pulsed Doppler signal recorded from the apical long-axis view as previously described. 13-15Cardiac output was equal to stroke volume multiplied by the heart rate. Great care was taken to ensure that the diameter was measured perpendicularly to the left ventricular outflow tract, trailing edge to leading edge just proximal to the aortic leaflets. Also, the pulsed Doppler sample was positioned so that the angle between the Doppler cursor and the left ventricular outflow tract was as close to 0” as possible (always <20”) and that the Doppler signal was truly representative of the left ventricular outflow tract flow as emphasized by Skjaerpe et a1.16Prosthetic valve areas were determined using the pressure half-time method proposed by Hatle et a1,17and by the continuity equation described by Nakatani et all* as follows: PVA = SVLvoT/VTI, where SVLVoT is stroke volume in the left ventricular outflow tract, and VT1 is the velocity-time integral of the mitral jet Doppler signal. Multiple determinations were done in each patient to ensure reproducibility of measurements; in patients with atria1 fibrillation (11 patients), 1444

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an effort was made to choose complexes most representative of the average heart rate and, particularly when applying the continuity equation, to evaluate left ventricular inflow and outflow complexes occurring at similar heart rates. Effective orifice areas for each size of Intact valve were calculated using data provided by the manufacturer (Medtronic Blood Systems), based on studies on a hydraulic model19 performed at the Georgia Institute of Technology by Dr. A.P. Yoganathan (used with permission). This model measures the transvalvular pressure gradients and pulsatile flow rates at different cardiac outputs, with a pulse rate of 70 beats/min and systole accounting for approximately 35% of the cardiac cycle. From this information, valve areas were calculated using the formula proposed by Yoganathan et all9 and a range of values was obtained for each prosthesis size, suggesting that effective valve areas increase as a function of the prevailing flow and pressure gradient. The mean value of each in vitro ares range (Table I) was used in this study for linear regression analysis against in vivo determinations. The in vitro areas were not available to the echocardiographer at the time of Dopp ler echocardiographic measurement of the effective orifice areas. Color-flow mapping was performed in multiple views to interrogate for the presence of mitral or aortic regurgitation. This was visually graded as l+ when the jet was very narrow and extended only a short distance beyond the valve, and 2+ when the jet was somewhat larger and longer but remained narrow at the level of valvular anulus and did not extend beyond the tip of the anterior mitral valve leaflet (aortic) or first third of the left atrium (mitral). Significant 3+ or 4+ mitral or aortic regurgitation was not observed in our patient population. Continuous variables are expressed as mean, standard deviation of the mean and range. Statistical analysis of the association of continuous variables was done using the Pearson linear correlation coefficient and standard error of the y estimate. Graphs were constructed with the corresponding linear regression equation.

lated by the pressure half-time method do not, however, correlate significantly with the in vitro valve areas (r = 0.15), with overestimations of the known in vitro area range occurring in 18 of 26 cases (6%) (Figure 2). There is similarly no correlation between the pressure half-time and continuity equation methods (r = 0.23). A significant correlation is observed between the peak and mean gradients (r = 0.85). Correlations of peak and mean gradients with areas are r = -0.60 and r = -0.64 for in vitro prosthetic areas, r = -0.59 and r = -0.63 for in vivo continuity equation-derived areas (Figure 3), and r = -0.05 and r = -0.11 for areas obtained with the pressure half-time method. Cardiac output does not correlate with body surface area (r = 0.28, p >0.05) in these patients and cardiac index is significantly lower in the 11 patients with atria1 fibrillation than in the 15 patients with normal sinus rhythm

For statistical inference, p values were obtained using the t test for the correlation coefficient and a p value <0.05 was considered significant. RESULTS

Table I lists the results for areas and gradients, as well as the incidence of valvular regurgitation according to prosthesis size. For the 26 patients, the peak gradient ranged from 5.4 to 21.0 mm Hg (mean 11.9 f 4.2) and the mean gradient from 2.3 to 12.7 mm Hg (mean 5.5 f 2.3). Valve areas ranged from 1.20 to 1.91 cm2 (mean 1.54 f 0.17) by the continuity equation and from 1.26 to 2.93 cm2 (mean 1.97 f 0.4) by the pressure half-time method. There is a good correlation (r = 0.82) between the mean in vitro area for each bioprosthesis size and the in vivo areas calculated by the continuity equation (Figure 1). The in vivo valve areas calcu-

FIGURE

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(2.4 f 0.4 vs 3.1 f 0.6 liters/min/m2, p
and to underestimate it in cases with mitral regurgitation. No patient in this study had significant (3+ or 4+) valvular regurgitation, but results could have been somewhat influenced by this factor in the patients with l+ or 2+ regurgitation. The pressure and flow data provided by the manufacturer suggest that there is not a unique area for each prosthesis size but rather that the effective orifice area tends to be augmented with increasing pressure and DISCUSSION Stroke volume as used in the continuity equation flow. This may be due to some leaflet inertia at low flow was measured in the left ventricular outflow tract for 2 rates. The relation given in Figure 1 might be improved reasons: (1) it is generally agreed that this is probably if it was possible to account more precisely for this facthe most reliable and easily accessible site for cardiac tor rather than using the mean in vitro areas for comparison. Nonetheless, the area values obtained in vivo output determinations by Doppler echocardiographylW; and (2) cardiac output determinations at the by the continuity equation method are within the range level of the mitral anulus are not possible in patients of the predicted in vitro area values in 24 of the 26 cases (92%), whereas there is a slight overestimation in one with mitral valve prostheses. This approach is theoretically not valid to measure mitral valve areas in cases patient (0.03 cm2) and an underestimation in another with valvular regurgitation, as it would tend to overesti- (0.12 cm2). The latter might be due to either deterioration of the valve in the 19 f 4 months since implantamate the valve area in cases with aortic regurgitation

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tion or to an error in methodology. Overall, the data suggest that the continuity equation, as applied in this study, is valid for assessing mitral valve bioprosthetic areas in the absence of significant valvular regurgitation. The valve area values calculated by the pressure half-time method do not, however, correlate with the in vitro values and, as demonstrated in Figure 2, this method grossly overestimates the effective valve area in most cases. Notwithstanding an unlikely error in the in vitro methodology, this fmding shows that the method is not valid to measure bioprosthetic valve areas and is consistent with recent reports7J8 The pressure half-time method was derived empirically in native valves and subsequent investigators have demonstrated that it may be influenced by many factors other than the valve area7,8 and may not be accurate in certain instances, such as after valvuloplasty.9 A possible explanation for the overestimation found in most of the patients with bioprosthesis might be the presence of some opening inertia of the prosthesis at the beginning of diastole, resulting in a higher initial gradient and a steeper deceleration slope once the valve has opened. Another might be the sudden change by outside intervention of a single parameter (effective valve orifice area) without enough time for adaptive adjustments to occur in the other parameters known to influence the half-time method. As in previous studies,5,6 Table I and Figure 3 show that peak and mean gradients are only grossly correlated to prosthesis size. In this context, it should be emphasized that pressure gradients are determined not only by valve orifice area but also by prevailing cardiac output, which varies from one individual to another depending on body size, hemodynamic status, cardiac rhythm and other factors. Hence, it would appear that valve area determinations are more accurate than gradient mea-

surements in assessing intrinsic prosthetic performance. Nonetheless, pressure gradients are also important because they reflect the increased stroke work created by the prosthesis and they may be unacceptably high, either because of pathologic stenosis, a state of high cardiac output or a mismatch between prosthesis size and body size.*O-** The lack of correlation in this study between cardiac output and body surface area probably reflects a heterogeneity of hemodynamic status in these patients. Using the relation derived between effective prosthesis area divided by cardiac output (cm*/liter/min) and mean gradient in the subgroup of patients with normal sinus rhythm and assuming a normal cardiac index of 3.3 liters/min/m2,23 one can nonetheless project (Figure 4) what the relation between effective prosthesis area indexed for body surface area and mean gradient would be in patients with normal cardiac output at rest as well as with 10 to 50% increases of stroke volume such as may occur during maximal upright exercise.24-26Since indexed effective prosthesis area can easily be calculated preoperatively from the in vitro calculated effective orifice area and the patient’s body surface area, this relation could be useful in selecting the type and size of prosthesis to be implanted to avoid valve prosthesis-patient mismatch. Further studies are necessary, however, to confirm some of these assumptions as well as their applicability to other types of prostheses. The performance of this prosthesis with regard to gradients and areas compares to that of other bioprostheses of similar sizes.1~3~4 Relatively high resting gradients (mean gradient >5 mm Hg) were found in 13 of 26 patients without evidence of pathologic prosthetic stenosis and are likely due to a mismatch between prosthesis size and body size. These results emphasize that bioprostheses tend to have relatively small orifice

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areas,27*28and the type of analysis proposed in Figure 4 could therefore be important to consider. Studies with exercise as well as studies relating prosthesis size to mortality, morbidity and functional class appear necessary to define the minimal requirements in terms of indexed prosthetic area. REFERENCES 1. Ryan T, Armstrong WF, Dillon JC, Fe&&urn H. Doppler echocardiographic evaluation of patients with porcine mitral valves. Am Heart J 1986;111:237244. 2. Panidis IP, Ross J, Mintz GS. Normal and abnormal prosthetic valve function as assessed by Doppler echocardiography. JACC 1986;8:317-326. 3. Reisner AS, Meltzer RS. Normal values of prosthetic valve Doppler echocardiographic parameters: a review. J Am Sot Echo 1988;1:201-210. 4. Sagar KB, Wann S, Paulsen WHJ, Romhilt DW. Doppler echocardiographic evaluation of Hancock and Bjork-Shiley prosthetic valves. JACC 1986;7.$81687. 5. Alam M, Rasman H, Lakier J, Kemp S, Khaja F, Hautamaki K, Magilligan DJ, Stein PD. Doppler and echocardiographic features of normal and dysfunctioning bioprosthetic valves. JACC 1987;10;851-858. 6. Fawzy ME, Halim M, Ziady G, Mercer E, Phillips R, Andaya W. Hemodynamic evaluation of porcine bioprostheses in the mitral position by Doppler echocardiography. Am J Cardiol 1987;59:643-646. 7. Loyd D, Ask P, Wranne B. Pressure half-time does not always predict mitral valve area wrrectly. J Am Sot Echo 1988;1:313-321. 8. Thomas JD, Weyman AE. Doppler mitral pressure half-time: a clinical tool in search of theoretical justification. JACC 1987;10:923-929. 9. Come PC, Riley MF, Diver DJ, Morgan JP, Safmn RD, McKay RG. Noninvasive assessment of mitral stenosis before and after percutaneous balloon mitral valvaloplasty. Am J Cardiol 1988,61:817-825. 10. Holen J, Aaslid R, Landmark K, Simonson S. Determination of pressure gradient in mitral stenosis with a noninvasive ultrasound Doppler technique. Acra Med &and

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21. Poliner LR, Demmer GJ, Lewis SE, Parkey RW, Blomqvist LG, WilIerson JT. Left ventricular performance in normal subjects: a comparison of the responses to exercise in the upright and supine positions. Circulation J980,62:528534. 26. Iskandrian AS, Hakki AJ, DePace NL, Manno B, Segal BL. Evaluation of left ventricular function by radionuclide angiography during exercise in normal subjects and in patients with chronic coronary heart disease. JACC 1983;1:15181529. 27. Nellessen U, Masuyama T, Appleton CP, Tye T, Popp RL. Mitral prosthesis malfunction. Comparative Doppler echocardiographic studies of mitral prostheses before and after replacement. Circulation 1989;79:330-336. 28. Goldrath N, Zimes R, Vered Z. Analysis of Doppler-obtained velocity curves in functional evaluation of mechanical prosthetic valves in the mitral and aortic positions. J Am Sot Echo 1988;1:21 l-225.