Durability/Wear Testing of Heart Valve Substitutes
`Helmut Reul, Klaus Potthast
`
`Helmholtz-Institute for Biomedical Engineering, Aachen, Germany
`
`Background and aims of the study: The current stan-
`dardsfor accelerated heart valve testing have consid-
`erable differences in test conditions. Another prob-
`lem arises from the fact that such test systems are not
`standardizedatall. It was shownearlierthat different
`test systems generate totally different valve loading,
`even if operating at standard conditions. The present
`study aimed to improvethis unsatisfactory situation
`and to develop a new concept whereactual loading of
`valves is measured either in vitro or in vivo under
`physiologic conditions and subsequently to repro-
`duce these conditions during accelerated testing.
`Methods:Integral loading forces at valve closure were
`measuredfor several valve types using a piezoelectric
`force ring within a real-time circulatory mock loop
`under physiologic conditions. This facilitated defini-
`tion of a physiologic loading range. Physiologic load-
`ing was subsequently reproducedina single-chamber
`accelerated test system. Working conditions obtained
`in termsof stroke, bypass flow and compliance served
`as design criteria for a new test chamber and a com-
`
`plete 12-chamberaccelerated testing system.
`Results: The integral loading obtained using the force
`ring showeda correlation with previous in vitro and
`in vivo results of strain-gauged valves. Loading
`forces for mechanical valves are about one order of
`magnitude higher than for bioprosthetic valves and
`are strongly related to cardiac output for both valve
`types. At physiologic loading, however, the differen-
`tial pressures across the valves are considerably
`below those given in FDA guidelines.
`Conclusions: This pilot study demonstrates that phys-
`iologic valve loading is reproducible over a wide
`range under appropriate testing conditions. It also
`showedthat,at the back-pressuresof the currentstan-
`dards, the loading forces during accelerated testing
`exceed the real-time loading forces by far and, thus,
`may provide unrealistically high valve loads. These
`initial findings indicate that amendments of the cur-
`rently valid standards maybe need to be accorded.
`
`The Journal of Heart Valve Disease 1998;7:151-157
`
`In 1996 we publisheda critical review on the state of
`the art of heart valve wearandfatigue testing as repre-
`sented by the current standards of ISO, FDA and CEN
`(1). These standards have considerable differences in
`test conditions as far as pressure differences across the
`test valves and total cycle numbers are concerned.
`Moreover, since the testing devices themselves are not
`standardized, different devices generate totally differ-
`ent loading conditions on the valves, even if the tests
`are carried out under the same standard. This study
`clearly demonstrated that, under the current stan-
`- dards, the conditions of actual, in vivo impact loading
`of a valve cannot be reproduced. Also, the twotest
`devices which were compared, although both operat-
`ing at the given standards conditions, generate totally
`different loading conditions. For one tester loading
`decreases with increasing cycle rate, while for the other
`it increases.Forall testers, time history of loading and
`
`Address for correspondence:
`Prof. Dr.-Ing. H. Reul, Helmholtz-Institute for Biomedical Engineer-
`ing, Pauwelsstr. 20, D-52074 Aachen, Germany
`
`pressure difference across the valve are a function of
`test frequency, compliance and otherfactors.
`To improve this unsatisfactory situation it was
`strongly suggested that the actual loading conditions
`of each valve type be measuredeither in animal mod-
`els in vivo or within a circulatory mock loop which pro-
`vides physiologic loading conditions;
`this loading
`could then be subsequently reproduced during accel-
`erated wear testing. The present study is set out to
`investigate this concept.
`
`Measurementsof valve loading
`under physiologic conditions
`
`Since most valves - and especially bioprostheses - are
`not suitable for the attachmentof strain gauges due to
`the lack of appropriate fixation points which are repre-
`sentative for valve loading, a piezoelectric force mea-
`surement
`ring (Kistler,
`type 906 1A, Winterthur,
`Switzerland) (Fig. 1, top right) was usedto assess load-
`ing at valve closure. This type of force transducer mea-
`sures the integral loading force acting on the valve.
`
`© Copyright by ICR Publishers 1998
`
`PAGE 1 OF 7
`
`WATERS TECHNOLOGIES CORPORATION
`
`EXHIBIT 1014
`
`WATERS TECHNOLOGIES CORPORATION
`EXHIBIT 1014
`
`PAGE 1 OF 7
`
`

`

`152 Durability/weartesting of heart valve substitutes
`H. Reul, K. Potthast
`
`J Heart Valve Dis
`Vol. 7. No. 2
`March 1998
`
`Integral Loading Forcein Vitro
`400|(Piezoelectric Force Ring)™~
`
`
`
`Strut Loading in Vitro
`\ eo
`=
`
`eu e
`
`a
`
`Strut Loading in Vivo
`
`seme Re
`
`;
`
`120
`
`
`
`Force/N—>
`
`ooaoOOo
`
`20
`
`
`
`( NF = 100 kHz )
`
`semiconductor
`pressure transducer
`
`
`inductive
`displacement
`transducer
`
`elektromagnetic
`ACA
`
`vibrator
`
`0
`
`500
`
`2000
`1500
`1000
`Left Ventricular dp/dt / mmHg/s >
`
`2500
`
`Figure 2: In vivo andin vitro loading of a 29 mm BSCC valve
`as a functionof left ventricular dp/dt before valve closure.
`
`methodsgive a good correlation with left ventricular
`dp/dt and, thus, can as well be correlated among each
`other.
`
`Measurements underaccelerated testing condi-
`tions
`
`Figure 1: Schematic of single-chamberaccelerated valvetest
`system with piezoelectric force ring in the upperright.
`
`After establishing the above correlation between inte-
`gral forces and left ventricular dp/dt the valve was
`inserted into a specifically developed single-chamber
`system foraccelerated valve testing (also shownin Fig.
`1).
`
`The next step wasto correlate these integral loading
`forces with previously obtained in vivo and in vitro
`results (2,3) measured bya strain-gauged 29 mm Bjérk-
`An electromagnetic vibrator (Ling Dynamics, type
`Shiley Convexo-Concave (BSCC) mechanical
`tilting
`409, Royston, UK) generates a sinusoidal flow through
`disk valve. For this purpose, the strain-gauged BSCC
`the test valve by compression and extension of a metal-
`valve was mounted onthe force ring and inserted into
`lic bellows. When the valveis closed, fluid from the
`the mitral position of a physiologic circulatory mock
`lower chamber flows to the upper chamber via an
`loop as previously described (4). Both, strut loads and
`adjustable bypass. This bypass serves simultaneously
`integral loading force were measured simultaneously.
`to control the pressure difference across the valve,
`For these measurements the left ventricular dp/dt
`which is measured upstream and downstream of the
`(according to the FDA guidelines averaged over the
`valve by two semiconductorpressure transducers with
`last 20 ms before valve closure) of the model ventricle
`a natural frequency of 100 kHz (Cobe Disposable
`wasvaried between 500 and 1500 mmH¢g/s by increas-
`Transducer, Lakewood, CO, USA). Test fluid is water
`ing the stroke volumeata fixed test rate of 70 per min,
`at room temperature.
`a mean aortic pressure of 100 mmHg,andasystolic
`Thetest rate was increased from 200 to 2000 per min
`duration of 15%.
`and the integral loading forces were measured. The
`The results are shown in Figure 2. The upper curve
`pressure difference across the valve was kept constant
`represents the force ring measurements which range
`at 120 mmHgforall test rates. The results are present-
`from 25 to 110 N integral loading force at the corre-
`ed in Figure3.
`sponding dp/dt values. The two lower curves repre-
`The shaded area represents the range of loading
`sent the strut loading forces obtained by strain gauges
`forces under all potential physiologic conditions as
`in vivo and in vitro for the same valve; these range
`obtained in vivo and in vitro and as depicted in Figure
`from 2 to 20 N strut loading force. The difference in
`2. The central curve showsthe measuredintegral load-
`measured forces between the two methods is obvious
`ing forces under accelerated testing conditions under
`andis related to the fact that the strain gauges measure
`observance of the FDA conditions for accelerated valve
`only loadingof a single strut caused bystrut deflection,
`testing such as full opening andclosing and 120 mmHg
`while the force ring measuresthe integral total loading
`pressure difference. The minimally adjustable loads
`force acting on the closed valve. Nevertheless, both
`
`
`
`PAGE 2 OF 7
`
`PAGE 2 OF 7
`
`

`

`Vol. 7. No. 2
`March 1998
`
`Durability/wear testing of heart valve substitutes
`H.Reul, K. Potthast
`
`153
`
` J Heart Valve Dis
`
`Oo
`
`600
`
`1000 1200 1400 1600 1800 2000 2200
`800
`Test rate / beats / min»
`
`*5
`
`200
`
`400
`
`Figure 3: Integral loading forces within the single chamber
`accelerated test system obtained for a 29 mm BSCCvalveat a
`back-pressure of 120 mmHg (FDA-condition).
`
`are represented by the lower curve. Higher loads can
`be easily generated by adjusting stroke and bypass
`throttle of the test system. These results clearly show
`that physiologic valve loading can be reproduced over
`a wide range under appropriate accelerated testing
`conditions. It is also evident that, under FDA condi-
`tions, only a single loading force which increases with
`highertest rates can be generatedfor a definedtestrate,
`whereastest conditions outside of the current standard
`in termsof pressure differencefacilitate the adjustment
`of a range of physiologic loading conditions at any test
`rate. Thus, valve loadingat resting conditions as well at
`exercise conditions can be adjusted, resulting ina much
`better simulation of the loading history of an implant-
`ed valve.
`
`.
`
`Valve loading forces for alternative valve types
`
`So far, all presented results are only valid for one single
`valve type and size, a 29 mm diameter BSCC tilting
`disk valve. Therefore, in order to develop a generally
`valid concept, it is necessary to extend the above find-
`ings to other valve types andsizes. For this purpose, a
`St. Jude Medical (SJM) 27 mm mechanical bileaflet
`valve and an Jonescu-Shiley 25 mm pericardial bio-
`prostheses were selected as a first step. Both valves
`were inserted into mitral position of our circulatory
`mockloop andintegral loading forces were measured
`underthe following experimental conditions (accord-
`ing to FDA guidelines for pulsatile flow valvetesting):
`
`Cardiac output: 2 1/min and 7 1/min;
`Test rate: 70 per min;
`Meanaortic pressure: 100 mmHg;
`Meanatrial pressure: 10 mmHg; and
`Systolic duration: 35%.
`
`PAGE 3 OF 7
`
`
`
`
`
`IntegralLoadingForce/N—>
`
`= rso
`
`= id°o
`
`= oo
`
`Bure
`
`40
`
`20
`
`Loading under FDA -Conditions
`
`Rangeof Physiologic Conditions
`
`Minimally Adjustable Loading
`
`Force/N—>
`
`LoO
`
`NMwOoOo
`
`Force/N—>
`
`= oO 0
`
`2L/min
`
`7L/min
`
`2L/min
`
`7L/min
`
`Figure 4: Integral loading forces for a 27 mm SJM mechanical
`valve and a 25 mm ISP bioprosthesis measured underreal-
`time conditions within a circulatory mock loop.
`This range of test conditions represents the normal
`physiologic range and can be usedfor the characteriza-
`tion of corresponding loading conditions. For each
`valve type the measured loading forces were averaged
`over 100 cycles.
`The results (Fig. 4) show that loading for the mechan-
`ical valve is about one order of magnitude higher than
`for the bioprosthetic valve, and loading increases with
`a factor of about two when cardiac output changes
`from 2 1/min to 7 1/min.
`
`Design of a new fatigue tester
`
`Theprincipal design criteria for a valve test chamberin
`terms of stroke, displaced volume, bypass flow and
`compliance were obtained by meansof the above-men-
`tioned single-chamber system (see Fig. 1). Based on
`these criteria a new test compartment for general use
`wasdesigned (shown schematically in Fig. 5). For the
`new design, the electromagnetic vibrator was replaced
`by a swashplate with adjustable stroke. Otherwise, the
`operating principle was the sameas already described
`above.
`Larger valves generate higher impact.loading forces
`than smaller ones. Therefore, since the new test com-
`partment is intended for use with all kinds of valve
`sizes, an adjustable compliance chamber wasadded for
`additional control of
`loading forces. The desired
`impact force can then be adjusted by varying the fol-
`lowing four parameters:
`
`- stroke of the swashplate;
`-
`test rate;
`- air compliance; and
`- bypass throttle adjustment.
`
`The FDA guidelines require that for all valves, test-
`ing should be conducted onthree of the largest, medi-
`um and smallest of each valve type. One equivalenttis-
`
`PAGE 3 OF 7
`
`

`

`154 Durability/wear testing of heart valve substitutes
`H.Reul, K. Potthast
`
`Test compartmentof HIA-FT2
`
`eee Observation Tube
`
`__— Test-Compartment
`
`— Air-Compliance
`
`_> Pressure Taps
`
`Compliance Chamber
`
`
`
`open to atmosphere ~
`
`Throttle —
`
`Testvalve
`Bypass
`
`——
`
`Piston Rod
`
`
`
`J Heart Valve Dis
`Vol. 7. No. 2
`
`March 1998
`
`into the new test compartment and the loading forces
`together with the pressure difference across the valves
`were measured within the fatigue tester at a test rate of
`1000 per min for the SJM and 600 per min for the ISP,
`respectively. The results (Figs. 8 and 9) show that for
`both valve types the peak loading forces are within the
`previously determined physiologic range (compare Fig.
`4). The differential pressures, however, are considerably
`below thepressures given in the FDA guidelines, which
`should be adjusted to 120 mmHgfor both valve types.
`An exampleillustrating this controversy for the SJM
`valve (Fig. 10) showsthat, in this case, the pressure dif-
`ference across the valve was adjusted to 120 mmHg by
`reducing the test chamber compliance. Test
`rate,
`bypass flow and stroke were kept constant. As can be
`clearly seen, valve loading forces exceed 50 N and,
`thus, are far above physiologic loading.
`
`Discussion
`
`
`
`——— Metal Bellow
`
`Figure 6: Top view of newly developed 12-chamberaccelerated
`test system (HIA-FT2).
`
`
`
`}-— ——— AdjustableStroke
`
`SwashPlate
`
`Figure 5: Schematic cross-section of newly developed test com-
`partment for accelerated valvetesting.
`
`sue annulus diameterof each type reference valve must
`be tested under identical conditions. This results in a
`total of 12 test valves. Accordingly, a 12-cylinder
`fatigue tester (HIA-FT2) has been designed and manu-
`factured (Fig. 6).
`Thefirst two of the above-listed parameters cannot
`be adjusted individually; they are the same for all 12
`test compartments. Thus, after the correct adjustment
`of one compartmentthe other 11 must be adjusted by
`variation of air compliance and bypassthrottle flow.
`An example (Fig. 7) shows where the peak loading
`forces for a 27 mm SJM valve have been varied by
`changing the air compliance at a constant test rate,
`bypassthrottle and stroke settings. Peak load decreas-
`es linearly with increasing compliance volume and can
`easily be adjusted for physiologic loading conditions.
`
`PAGE 4 OF 7
`
`This pilot study demonstrates that physiologic valve
`loading can be reproduced under appropriate acceler-
`ated testing conditions. However, testing under these
`conditions requires special testing compartments with
`an increased number of control parameters for the
`adjustmentof physiologic loads. It also requires a two-
`step testing approach:first, physiologic loading forces
`have to be determined within a real-time circulatory
`mock loop; and second, this loading has to be repro-
`duced within the accelerated tester. For propertransfer
`
`Inafinal step a 27 mm mechanical SJM valve and a 25 of physiologic real-time loading forces to accelerated
`testing the test valves have to be mounted within a cal-
`mun lonescu-Shiley pericardial valve (ISP) were inserted
`
`Verification of concept within new fatigue tester
`
`PAGE 4 OF 7
`
`

`

`| Heart Valve Dis
`Vol. 7. No. 2
`March 1998
`
`
`
`
`
`
`Force/N-=
`
`
`
`70
`
`100
`90
`80
`Compliance Volume / ml—
`
`110
`
`Physiologic Range
`according to FDA-Guidelines
`
`Figure 7: Integral loading force as a function of air compliance
`volume measured for a 27 mm SJM valve within the new test
`compartment at constant test rate (1000 per min), bypass
`throttle andstroke settings.
`
`ibration compartment, equipped with a force ring
`transducer. Once physiologic loading is adjusted and
`the parametersettings in terms of stroke volume, com-
`pliance volume and bypass flow are obtained,
`the
`valves are mounted within the geometrically similar
`final test compartments and the continuously mea-
`sured pressure difference serves as the single control
`parameterfor long-term studies.
`It could also be shownthatat the back-pressures of
`the current standards the loading forces during accel-
`erated testing exceed thereal-time loading forces by far
`and thus mayprovide unrealistically high valveloads.
`These initial findings indicate that corresponding
`amendments of the currently valid standards may be
`necessary.
`
`25 mm ISP
`
`Physiologic Range
`
`Force/N—
`
`Ap/mmHg—>
`
`0,0
`
`01
`
`03
`0,2
`Time/s—>
`
`04
`
`05
`
`Figure 9: Integral loading force and pressuredifference across
`valve for a 25 mm ISP bioprosthesis at a test rate of 600 per
`min,
`
`PAGE 5 OF 7
`
`Durability/wear testing of heart valve substitutes
`H.Reul, K. Potthast
`
`155
`
`
`
`Force/N—»
`
`27 mm SJM
`
`4‘
`Physiologic Range
`
`Ap/mmHg—> Time /s —>
`
`Figure 8: Integral loading force and pressure difference across
`valve for a 27 mm SJM valveata test rate of 1000 per min.
`
`Durability testing of bioprosthetic heart valves
`requires some additional discussion. The FDA and ISO
`guidelines for the weartesting of tissue valves are very
`similar to those for mechanical valves. Valves are to be
`tested for an equivalent of five years at a peak back-
`pressure of at least 90 mmHgfor aortic and 120 mmHg
`for mitral valves. The valves should open and close
`completely. For stentless valves, the aortic wall should
`be modeled by a compliant tube.
`Unlike mechanical valves, tissue valves may show
`severe damageas a result of the wear test and may
`becomedysfunctional before completion of the experi-
`ment. Possible damage to tissue valves include tissue
`tear, holes, delamination, abrasion, prolapse and stent
`fracture (5,8,9). Table I lists commonly observed dam-
`age in tissue valves, their cause, andcritical test para-
`meters that influence the damage. Wear in mechanical
`valvesis also includedasa reference.
`Tissue tear is typically caused by hightensile stress-
`es. Finite element modeling (6) indicates that the high-
`est tensile stresses are in the region of the commissures
`150
`
`27mm SJM
`
`8
`
`Force/N—>
`
`is
`
`Ap/mmHg—> -100
`
`0,00
`
`0,04
`Time /s —>
`
`0,08
`
`Figure 10: Integral loading force and pressure difference across
`valve for a 27 mm SJM valveat a test rate of 1000 per min,
`but at a preselected pressure difference of 120 mmHg (FDA-
`condition) across the valve.
`
`PAGE 5 OF 7
`
`

`

`156 Durability/weartesting of heart valve substitutes
`H.Reul, K. Potthast
`
`J Heart Valve Dis
`Vol. 7. No. 2
`March 1998
`
`
`
`Table I: Wear and damageobservedin cardiac valve prostheses during accelerated weartesting.
`
`Valve type
`Wear/damage
`Cause
`Critical physical
`Critical test
`
`parameters
`parameter
`
`Mechanical _
`Sliding wear
`Sliding motion
`Range of motion
`Full opening and heart
`
`valve
`closing
`
`Impact wear
`Inertia at impact
`Impactforce
`Ap, dp/dt, impact force
`
`
`
`
`Stented tissue valve__Tissuetearat Hightensile stresses Pressure force at Ap
`
`commissure closure
`
`
`Contact between
`Tissue deflection
`Tissue abrasion at
`commissure
`during valve opening
`tissue and stent during
`
`during valve opening
`Tissue deflection
`Contact between
`Abrasion at the base
`Ap
`
`at closure
`of the stent
`. leaflets and stent
`
`Maximum opening
`
`Delamination of
`Tissue deflection
`High bending
`tissue
`stresses, tissue
`during cardiac cycle
`
`buckling
`
`Ap, maximum opening
`
`Stentless tissue valve
`
`Pressure force at
`Tissue tear at
`AP, Piny aortic wall
`commissure
`compliance
`closure, aortic wall
`
`motion
`
`Hightensile stresses
`
`Abrasion at the base
`Contact between
`AP, Pim aortic wall
`leaflets and inflow
`closure, aortic wall
`compliance
`of cusp
`
`rim
`motion
`
`Leaflet deflection at
`
`Tissue deflection
`Delamination of
`High bendingstresses,
`AP, Piny Maximum
`tissue
`opening,aortic wall
`during cardiac cycle
`tissue buckling
`compliance
`Axial and radial
`Tearin aortic wall
`Hightensile stresses
`AP, Piny
`wall compliance
`pressure loading
`AP: back-pressureacrosstheleaflets; p,,,: transmural pressure acrosstheaortic wall.
`
`in response to the pressure loading at closure. It is
`therefore important to control the closing pressure
`accurately. In the case of stentless valves, the stresses in
`the leaflets may be redistributed by the flexible aortic
`wall which mayactas a shock absorber. Proper model-
`ing of the aortic wall compliance and control of the
`transmural pressure gradients should therefore be con-
`sidered whentesting stentless valves.
`tissue
`Tissue abrasion is caused by rubbing of
`against the stent, stent coveror tissue reinforcement.It
`most likely occurs close to the commissuresor at the
`base of the cusps where excursion ofthe tissue is max-
`imum during the cardiac cycle. Maximum deformation
`of the leaflets in the cusp region occurs at maximum
`back-pressure, while maximum deformation of the
`leaflets at the commissure occurs during full opening.
`Thus, control of both peak back-pressure and maxi-
`mum leaflet opening is important when testing for
`leaflet abrasion. It should be noted that the high cycle
`
`rate in accelerated weartesters may notcreate the same
`leaflet deflection or stresses as in a real-time pulse
`duplicator. Vesely et al. (10) pointed out that porcine
`tissue is viscoelastic,i.e. the stress-strain relationship is
`time-dependent.
`Delamination in the tissue is due to internal shear
`stresses. Vesely and Boughner (7) have shown that
`porcine tissue loses someofits naturalability to shear
`when tanned. As a result, bending of tanned tissue
`leads to high internal shear stresses, and potential
`buckling of the tissue (8). Delamination is typically
`‘observed in the hinge area of porcine valves where the
`tissueis relatively thick.
`Asindicated in TableI, critical test parameters for the
`testing of tissue valves include the maximum back-
`pressure and the maximum valve opening. In some
`commercially available durability testers,
`the back-
`pressureis controlled by a throttle in the flow loop that
`adjusts the bypass flow during closure. As pointed out
`
`PAGE 6 OF 7
`
`PAGE 6 OF 7
`
`

`

`Durability/wear testing of heart valve substitutes
`H. Reul, K. Potthast
`
`157
`
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`
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`6. Krucinski S, Vesely I, Dokainish MA, Campbell G.
`Numerical simulationof leaflet flexure in bioprosthet-
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`7. Vesely I, Boughner DR. Analysis of the bending
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`and
`shear
`measurements.
`J
`Biomech
`1989;22(6/7):S655-S671
`8. Vesely I., Boughner DR, Song T, Tissue buckling as a
`mechanism of bioprosthetic valve failure. Ann Thorac
`Surg 1988;46:5302-S308
`9. Thiene G, Bortolotti U, Valente M,et al. Modeoffail-
`ure of the Hancock pericardial xenograft. Am J Car-
`diovasc Surg 1989;63:S129-S133
`10. Vesely I, Boughner DR, Leeson-Dietrich J. Biopros-
`thetic valve tissue viscoelasticity:
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`1995;60(2):S379-S382
`
`] Heart Valve Dis
`Vol. 7. No. 2
`March 1998
`
`earlier, tissue valves may experience notable damage
`during the experiments which often leads to increased
`backflow leakage. To maintain the required peak back-
`pressure, the bypass resistance has to be increased to
`compensate for leakage through the test valves.It is
`therefore desirable to have individualthrottles for each
`valve in order to maintain the same back-pressure for
`all valves throughoutthe duration of the experiment.
`The second importanttest parameteris the degreeof
`opening of tissue valves. According to the regulatory
`guidelines, full opening should be achieved in the
`durability tester.
`‘Full opening’
`is somewhat
`ill-
`defined, as the degree of opening is a function of the
`flow rate. At Baxter Laboratories, for instance, the
`valve is first tested in a pulse duplicator at a cardiac
`output of 51/min. The degree of opening during peak
`forward flow is recorded with a video camera posi-
`tioned along the axis of the valve. When the valveis
`placedin the durability tester, the stroke of the actuator
`is adjusted to reproduce the same degree of opening as
`in the pulse duplicator. Video images of the valve in
`both systemsare used to verify proper tuning.
`In conclusion, the above discussion addresses some
`of the basic considerations for the testing of tissue
`valves. Because of the complex biomechanical proper-
`ties of tissue, it is unclear if and howthe accelerated
`cycling rate is modifying the stress andstrain distribu-
`tion in the valves (10). Furthermore, biochemical
`degradation andin situ host response are not consid-
`ered. Biochemical degradation or host overgrowth may
`modify the tissue properties and alter the motion and
`stress distribution in the tissue. Results of accelerated
`weartesting of tissue valves should therefore be
`reviewed in the context of the limitations of the test.
`Additional knowledge aboutthe viscoelastic behavior
`of fresh and aged tissue will be necessary to further
`refine the accelerated wear testing of tissue valves.
`
`Acknowledgments
`The authors thank Stefan G. Schreck PhD, Director,
`Engineering Research and Standards at Baxter Health-
`care Corporation, Irvine, CA and Carlos Rios from the
`same group for their most valuable contribution to the
`discussion on specific aspects for accelerated testing of
`bioprostheses.
`
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