Patent Description:
For ex-vivo assessment of the viability of a donor heart ventricular pressure-volume recording can provide an indication of the pumping behaviour of a donor heart. Apparatus for performing measurements on a heart ex-vivo are typically used for studies of the human or animal heart in scientific research.

For instance <NPL>), discloses in-vitro measurements on a beating heart involving use of a preload chamber for generating a preload, a transapically placed PV catheter extending through an incision, a roller pump and a centrifugal pump for controlling afterload.

In <NPL>, measurements on a pig heart are described for which the left atrium was opened between the pulmonary veins, and an artificial, Y-shaped valve apparatus, constructed from stiff plastic tubes was inserted. An artificial valve was placed in the aortic branch and one in the atrial branch of the apparatus, allowing only unidirectional flow in each branch. A flexible latex balloon was tied over the tip of the ventricular branch and inserted into the left ventricle through the mitral valve of the isolated heart. A plastic tube, connected to the aortic branch of the apparatus, was elevated <NUM> above the heart before entering an open (atrial) reservoir. With a flow rate of <NUM>/min through this part of the tube system the resistance was calculated to be <NUM>,<NUM> dynes. cm-<NUM>, which is described to be comparable with the systemic vascular resistance measured in the donor pigs before harvesting of the heart. Another tube was connected the bottom of the atrial reservoir with the atrial branch of the valve apparatus. Balloon, tubes, and reservoir were filled with <NUM>% NaCl at <NUM>. The reservoir was placed at the level necessary to create a filling pressure of <NUM> Hg in the balloon inserted into the left ventricle of the non-beating heart. A second flow probe was implanted to measure the flow rate through the aortic branch, i.e. cardiac output (CO).

In clinical practice however, reliable indications for predicting the viability of a donor heart need to be obtained with minimal interference and in particular minimal damage to the heart and minimal risk of contamination of the heart.

It is further noted that an apparatus according to the pre-characterizing portion of the appended independent claim <NUM> is disclosed in "<NPL>".

It is an object of the present invention to provide an apparatus and a method for measuring indications for predicting the viability of a donor heart with minimal interference and in particular minimal damage to the heart and minimal risk of contamination of the heart.

According to the invention, this object is achieved by providing an apparatus according to the appended independent claim <NUM>. For achieving this object, the invention also provides a method according to the appended independent claim <NUM>.

In view of the simple design with few parts of the provided or used apparatus, sterility of all surfaces to which the heart and fluids flowing to the heart can be exposed in a reliable and simple manner.

Particular elaborations and embodiments of the invention are set forth in the dependent claims.

Further features, effects and details of the invention appear from the detailed description and the drawings.

In <FIG>, an example of an apparatus <NUM> according to the invention for ex-vivo measurement of performance of a donor heart <NUM> is shown. The donor heart is shown in cross-section along a plane through the left ventricle <NUM>, the aortic valve <NUM> and the mitral valve <NUM>. The heart <NUM> is held in a heart holder <NUM> which is in this example a flexible support <NUM> in a receptacle <NUM> with a bottom part <NUM> and a lid <NUM>. The lid <NUM> hermetically seals off the receptacle <NUM> when in closed condition coupled to the bottom part <NUM> of the receptacle <NUM>.

For perfusing the heart <NUM>, a perfusion supply conduit <NUM> extends into and through the receptacle <NUM> and is connected to an aortic branch by a connector <NUM>. The perfusate introduced into the aorta flows as coronary flow from the aorta through the coronary arteries (not shown) of the heart and drains into the reservoir <NUM> via the right atrium <NUM> and optionally via the right ventricle <NUM>. For discharging fluid drained from the heart, a perfusion discharge conduit <NUM> is connected to the receptacle <NUM>. A perfusion pump unit <NUM> and an oxygenator <NUM> are connected in fluid communication between the perfusion supply conduit <NUM> and the perfusion discharge conduit <NUM>. An oxygen supply line <NUM> is connected to the oxygenator <NUM>. For controlling the temperature of the heart <NUM>, a thermostatic control unit <NUM> communicates with the oxygenator <NUM> via conduits <NUM>, <NUM>, through which a heat transfer medium such as water is circulated.

If a heart <NUM> is heated up from a cold non-beating condition, the heart will generate its own rhythm, while defibrillating actions may be required. The pace of the heart may also be adjusted or conducted by an artificial pacemaker.

A bag <NUM> is arranged in the left ventricle <NUM> of the heart (see also <FIG>). The bag <NUM> bounds a bag interior space <NUM> communicating with an interior space <NUM> (<FIG>) of a fluid tight, compressible and expandable container <NUM> on the atrium side of the mitral valve <NUM>, so that fluid can flow from the interior space <NUM> of the bag <NUM> into the interior space <NUM> of the container <NUM> and back. Furthermore, a sensor <NUM> for measuring compression and expansion of the container <NUM> is provided. In this example, the sensor <NUM> is coupled to the container <NUM> via Bowden cables <NUM>. Flow from the cavity <NUM> of the left ventricle through the aortic valve <NUM> to the aorta is blocked by the bag <NUM>, and so the passage through the mitral valve <NUM> is the only inlet and outlet of the ventricular cavity <NUM>. This allows inflow and outflow to be detected by volume changes of a single vessel <NUM> in fluid communication with the bag <NUM>.

The fluid in the bag <NUM> and in the container <NUM> may be gaseous (e.g. air) or liquid. An advantage of a gaseous fluid is that the combination of the bag <NUM> and of the container <NUM> is light, that the fluid itself generates very little flow resistance and that the fluid has very little thermal capacity, so that thermal control of the fluid is of relatively little importance. An advantage of providing a liquid as the fluid in the bag <NUM> and in the container <NUM> is that it is not compressible and, compared with a gaseous fluid, its flow resistance and specific mass is more similar to those of blood with regard to the way it loads the left ventricular <NUM>.

The bag <NUM> has a bag wall which rests against the left ventricle wall, so that, at its maximum volume during the measurements, the bag wall is substantially free from elastic stretch in a plane of the bag wall and subjected to very little strain in the plane of the bag. Thus, pressure and volume in the interior <NUM> of the bag <NUM> and in the interior <NUM> of the container <NUM> are not significantly influenced by elastic stretching of the wall of the bag <NUM>. The bag <NUM> is shaped to be larger than the shape of an interior of the left ventricle by a small margin only and of a shape similar to the shape of the interior of the left ventricle <NUM>, to limit bulging towards the aortic valve <NUM>. In order of increasing preference, the oversize in any direction is preferably not more than <NUM>, <NUM> resp.

The heart holder <NUM> is arranged in the receptacle <NUM>, which also encloses the heart <NUM>, the bag <NUM> and the container <NUM>. The sensor <NUM> is composed of a motor <NUM> with an encoder and a control unit <NUM> and also forms an actuator. The sensor and the actuator may also be provided as separate items, for instance by providing a force sensor arranged for sensing forces exerted via a transfer mechanism between the actuator and the container.

The sensor <NUM> is arranged outside of the enclosure <NUM>. The Bowden cables <NUM> form a motion transfer mechanism extending from the container <NUM> to the actuator and sensor <NUM> for transferring motion from the container <NUM> to the sensor <NUM> and from the actuator <NUM> to the container <NUM>. Thus, the actuator and sensor <NUM> can be arranged outside of the interior of the enclosure <NUM> of which sterility needs to be ensured. This facilitates sterilization and avoids the need of subjecting the actuator and/or sensor to sterilization treatments. Provision of alternative motion transfer mechanisms, such as a belt drive, a chain drive, a toothed rack drive, a (preferably non self-braking) spindle drive, a hydraulic drive and a lever with push and/or pull rods, are conceivable as well.

The Bowden cables <NUM> form a simple and flexible motion transfer mechanism that allows easy handling of the bag <NUM> and the container <NUM> when coupled to the motion transfer mechanism. By coupling the container <NUM> to the motion transfer mechanism before insertion of the bag <NUM> into the left ventricle <NUM>, the need of manipulation of and around the exposed heart <NUM> is reduced.

Core cables <NUM> (i.e. internal cables) of the Bowden cables <NUM> are coupled to a support <NUM> to which the a first end of the container <NUM> and an open end of the bag <NUM> are mounted and which is essentially stationary relative to the area of the mitral valve <NUM> via which the bag <NUM> projects into the left ventricle <NUM>. A second end of the container <NUM> opposite to the first end of the container <NUM> is coupled to an actuating member <NUM> to which distal ends of the outer cables <NUM> of the Bowden cables <NUM> are coupled. Proximal ends of the outer cables <NUM> of the Bowden cables <NUM> are coupled to a Bowden cable abutment <NUM> which is in a fixed position relative to an axis of rotation of a shaft <NUM> and a drum <NUM> mounted to that shaft <NUM>. If the core cables <NUM> are pulled by rotation of the drum <NUM>, distal ends of the corresponding core cables <NUM> and outer cables <NUM> move towards each other. This causes the support <NUM> and the actuating member <NUM> to move towards each other, so that the opposite ends of the container, which are coupled to the support <NUM> and the actuating member <NUM>, are moved towards each other and the container <NUM> is compressed. Conversely, if the core cables <NUM> are veered out by rotation of the drum <NUM> in an opposite sense, the container <NUM> expands. Veering out of the cables <NUM> preferably occurs passively or while a braking force exerted by the motor <NUM> is used to generate and control an afterload against which the heart <NUM> has to pump.

If, as in the present example, the container <NUM> is arranged for expansion in substantially one direction only and compression in an opposite direction only, the volume changes of the container <NUM> can be detected in a simple manner by detecting the expansion and compression in these, mutually opposite, directions. Other manners of detecting changes of the fluid volume in the container are conceivable as well but may be outside the scope of the appended claims, for instance detection of a fluid level in a rigid container or detection of a fluid volume in the rigid container, for instance by means of a level of an indicator liquid separated from the fluid flowing into and out of the ventricle by a membrane.

The container <NUM> is in the form of a bellows. This allows the expandability and compressibility in essentially one direction only to be achieved in a simple manner. Furthermore, sterility of the interior of a bellows can be reliably ensured.

Using an apparatus according to the invention, cardiac output can be measured in a simple manner by measuring expansion and contraction of the container <NUM>. Because of the simple design of the apparatus, with few parts, sterility of all surfaces to which the heart <NUM> and fluids flowing to the heart <NUM> can be exposed, can be ensured in a reliable and simple manner.

The motor <NUM> serves for exerting a compression force on the container <NUM> and the motor control unit <NUM> serves for controlling the compression. This allows to periodically exert a preload urging fluid from the container <NUM> into the bag <NUM> in the left ventricle <NUM>, which simulates a pressure at which, in-vivo, blood is pressed from the left atrium into the left ventricle <NUM> during the last part of diastole. The preload may be controlled by controlling the compression force and thereby a preload pressure, by controlling compression displacement and/or motion and thereby preload fluid displacement and/or flow, or by a combination of force and displacement and/or motion in accordance with a predetermined constant or varying relationship.

In the present example the functionality of the left ventricle is measured. In a similar manner, the functionality of the right ventricle <NUM> may be tested.

The motor <NUM> also forms a transducer for outputting a signal in response to an expansion of the container19, as a result of pumping action by the heart, to the control unit <NUM>. The control unit <NUM> is arranged for controlling and registering a counterforce exerted by the motor <NUM> onto the container <NUM> against the expansion of the container <NUM>. Such a counterforce simulates an afterload encountered in-vivo during systole as the left ventricle <NUM> contracts and causes blood to be expelled into the aorta. Thus, the combination <NUM> of the motor <NUM> and the controller <NUM> operate as a sensor and as an actuator.

The counterforce can be varied during each cycle to simulate the resistance and elasticity of the cardiovascular system and the inertia of blood in that cardiovascular system. The counterforce may be controlled as a function of fluid displacement and/or flow as sensed in the form of sensed expansion and/or expanding motion of the container.

As is shown in <FIG> and in the diagram shown in <FIG>, preloads and afterloads during the ex-vivo heart validation test are generated in a simple manner by influencing expansion and compression of the container <NUM>. This could for instance, in part, be achieved by providing the container and/or material in which it is embedded with suitable mechanical properties. However, in addition, or substantially as an alternative, an active control of preload and afterload, is advantageous to also allow simulation of the in-vivo effects of the closing of the aortic valve and the inertia of the blood.

The container <NUM> and the bag <NUM> in the ventricular cavity <NUM> bound a common, hermetically enclosed volume. When the heart <NUM> is contracting (<FIG>) pressure in the bag <NUM> in the left ventricle <NUM> increases (see dial P) and fluid is expelled from the bag <NUM> into the container <NUM> which causes the container volume to increase while pressure remains substantially constant (<FIG>).

The pressure in the container <NUM> can be determined by measuring a force exerted by the expanding container <NUM>, which force can also be influenced by a resistance to which expansion of the container <NUM> is subjected. The relationship between the pressure in the container <NUM>, the exerted force and a surface area of the container <NUM> facing in a direction opposite to the direction in which the force is exerted can be expressed as: <MAT> in which P = pressure, F = exerted force and A = surface area facing in a direction opposite to the direction in which the force is exerted.

The relationship between the fluid volume displaced into the container <NUM>, the distance of displacement of a container wall and a surface area of the displaced container wall facing in a direction opposite to the direction of wall displacement can be expressed as: <MAT> in which ΔV = displaced volume, D = distance of displacement of a container wall and A = surface area of the displaced container wall facing in a direction opposite to the direction of displacement.

To measure the distance of displacement, for instance an encoder coupled to the shaft <NUM> of the motor <NUM> can measure the angular movement of the shaft <NUM> and of drums <NUM> on the shaft <NUM>, around which drums <NUM> ends of core cables <NUM> of the Bowden cable <NUM> are wound. This encoder generates a number of pulses per turn. As the container <NUM> is filled, it expands in one direction. This expansion is converted into a turning motion of the motor shaft <NUM>. The rotation of the motor shaft <NUM>, and accordingly the number of pulses generated by the encoder, is therefore linearly related to the volume of fluid displaced into the container <NUM>.

The force the motor <NUM> exerts, is linearly proportional to the electrical current fed to or generated by the motor <NUM>, so that, by controlling the electric current to and from the motor <NUM>, pressure in the container <NUM> is controlled and monitored. An additional pressure sensor in the ventricular cavity <NUM> (inside or outside of the bag <NUM>) or in the container <NUM> can increase the accuracy of the pressure measurement.

After at least most of the displacement of fluid into the container <NUM>, the exerted pressure is reduced and flow of fluid into the container <NUM> comes to a standstill (<FIG>).

Subsequently, the ventricular wall relaxes and fluid flows back into the bag <NUM> in the left ventricle <NUM>. A force exerted onto the container <NUM> results in a preload pressure in the bag <NUM> in the left ventricle (<FIG>), which simulates preload from the left atrium during diastole.

Thus, ventricular pressure and ventricular volume can be measured through a heartbeat cycle and represented in a pressure/volume (PV) loop. An example of such a PV-loop is schematically shown in <FIG> in which the phases illustrated in <FIG> are indicated by corresponding reference numbers in circles.

The PV curve plots the ventricle pressure (Y-axis) versus ventricle volume (X-axis). Starting at the bottom left corner, opening of the mitral valve <NUM> is simulated and an inflow of blood into the bag <NUM> in the left ventricle <NUM> causes the left ventricle volume to increase, while pressure in the left ventricle <NUM> increases only slightly as a result of simulation of preload from the atrium (the phase shown in <FIG>).

When the left ventricle <NUM> has been filled, the ventricles contract, so that pressure increases. Resistance exerted onto the container <NUM> simulates closing of the mitral valve and counter pressure from the aorta (the phase shown in <FIG>).

In-vivo, the increasing pressure in the left ventricle would subsequently cause the aortic valve <NUM> to open when this pressure has exceeded the pressure in the aorta, and would cause an outflow of blood into the aorta. This is simulated by allowing the container <NUM> to expand while the pressure still rises simulating the increasing flow resistance and elasticity of the cardiovascular system and inertia of the blood displaced by the pumping action of the heart (the phase shown in <FIG>).

Next, the heart relaxes, so the pressure decreases. The decrease of pressure with no or very little increase of volume in the left ventricle <NUM> (the phase illustrated in <FIG>) simulates the in-vivo stage starting when the aortic valve <NUM> closes up to the moment when pressure in the left ventricle <NUM> becomes lower than the pressure in the left atrium so that the mitral valve <NUM> opens and the left ventricle <NUM> fills with blood.

In the proposed concept the pressures, volumes, valve openings and closings can be simulated by software. The control of exerted forces can be based on controlled pressure or based on controlled flow rate. Control of pressure can be based on measured flow rate and controlled flow rate can be based on measured pressure. In vivo, the inflow and the outflow are related to the ventrical pressure in accordance with flow characteristics, in particular the flow resistance and elasticity of the cardiovascular system and inertia of the displaced blood. The flow characteristics of the outflow preferably simulate in vivo flow characteristics of outflow into the aorta and all successive vessels and organs and the flow characteristics of the inflow preferably simulate in vivo flow characteristics of flow from left atrium to left ventricle <NUM>.

The combination of flow resistance, elastic compliance and inertia is defined as the impedance. The impedance may be adjustable and also variation of impedance over a cycle may be provided for and may also be adjustable. In particular, preload may gradually be increased for testing stress resistance. A viable heart will typically increase output (heterometric autoregulation) in response to an increasing preload while a heart with poor viability will typically return a decreased output if preload is gradually increased. Also muscle cell contraction can be tested at different afterload levels.

For the outflow, after the aorta valve opens, the flow rate is related to the pressure in accordance with the impedance of the aorta and downstream vessels and organs: <MAT> in which P = pressure, Q = flow rate and Iaorta = the fluidic impedance of the aorta and downstream vessels and organs.

To generate a simulation of in-vivo pressure P, the measured flow rate is multiplied by a predetermined impedance, the calculated pressure is multiplied by the surface area in the direction of displacement in accordance with equation (<NUM>) and the motor control <NUM> controls the motor <NUM> to exert the calculated force F. The force is controlled by the electrical current applied to the motor <NUM>.

The flow rate can be measured by sensing the volume change of the container <NUM> over time. In formula: <MAT> in which Q = flow rate, ΔVContainer = displaced container volume and ΔT = time over which displacement has been measured.

The change of volume in the container <NUM> can be measured by registering encoder pulses that are generated in accordance with angular movement of the shaft <NUM> and the drums <NUM> on the shaft around which ends of the core cables <NUM> of the Bowden cable <NUM> are wound. The time T between two pulses provides the ΔT for a volume change ΔV associated to the displacement of the container wall from one pulse to the next pulse. From these values the average flow rate Q in the time period between the two pulses in accordance with equation (<NUM>) can be calculated. When this flow rate Q is multiplied with the predetermined impedance I, the pressure P to be exerted and, via equation (<NUM>) the force F the motor <NUM> should exert on the core cables <NUM> connected to the container <NUM>, can be calculated. From the calculated force F, the electrical current to be applied to the motor <NUM> can be determined, for instance by the motor control <NUM>.

Because the calculated force F to be exerted trails the measured flow rate Q and because reaction by the motor <NUM> and the motor control <NUM> may take a response time, the exerted force F lags the flow rate Q on the basis of which it is determined. By arranging the system such that, at a given flow rate Q, encoder pulses are generated at a high frequency and carrying out the successive calculations of new values of force F at a high frequency as well, the time lag between flow rate measurement and exertion of the associated force can be reduced and a smooth adaptation of pressure P to flow rate Q can be achieved. The calculated force F may also be determined on the basis of a predicted value of the flow rate Q at the time the force F is to be applied. The prediction of the flow rate Q can for instance be made by Kalman filter calculations applied to previous predicted and measured values of the flow rate Q.

After the apparatus has been controlled to run during outflow into the aorta while the force F is controlled in accordance with the measured flow rate Q and virtual impedance of the aorta Iaorta (<FIG>), displacement of the container wall stops because the left ventricle has reached its contracted state and pressure ebbs away as a result of an absence of flow Q (<FIG>) until an initial preload pressure is reached. As the left ventricle <NUM> relaxes, the preload pressure is controlled to increase as inward flow is detected (<FIG>). This simulates the in-vivo opening of the mitral valve <NUM>.

In all phases of the heartbeat, the force to be exerted can be calculated by equations (<NUM>) and (<NUM>), each phase having a different impedance. However, the apparatus can also be controlled in a flow-controlled mode. In the flow controlled mode of operation, the pressure P will be measured and the flow rate Q is controlled in accordance with the measured pressure P, the pressure - flow rate equation (<NUM>), the flow rate displaced volume equation (<NUM>) and the relation between volume and displacement of equation (<NUM>).

It is also possible to switch between pressure controlled mode and a flow rate controlled mode during each heart beat cycle. For instance for operating in the pressure control mode during systole when the left ventricle actively pumps against a predetermined passive resistance (afterload) and for operating in a flow rate control mode during diastole in which the left ventricle is substantially passive and a predetermined pressure (preload) is applied to the left ventricle.

During operation, also the stroke volume of the left ventricle can be measured by measuring the total displacement of the container wall sensed by the encoder of the motor <NUM> and applying equation (<NUM>). Alternatively, the left ventricle volume can also be measured by measuring the total fluid volume during filling and subtracting the volume of the container <NUM> from the total volume.

Claim 1:
An apparatus (<NUM>) for ex-vivo measurement of performance of a donor heart, comprising:
a heart holder (<NUM>) in a receptacle (<NUM>) for holding a human donor heart (<NUM>);
a perfusion fluid supply conduit (<NUM>) extending into the receptacle with a connector (<NUM>) for connection to the heart;
a perfusion discharge conduit (<NUM>) communicating with the interior space of the receptacle for collection of perfusion fluid drained from the heart;
a perfusion pump (<NUM>) and an oxygenator (<NUM>) connected between the perfusion supply conduit and the perfusion discharge conduit;
a bag (<NUM>) for placement into a left ventricle (<NUM>) of the heart, the bag having a bag interior space (<NUM>); and
a fluid tight, compressible and expandable container (<NUM>) having a container interior space (<NUM>) communicating with the bag interior space (<NUM>):
characterized by:
a sensor (<NUM>) for measuring compression and expansion of the container (<NUM>);
an actuator (<NUM>, <NUM>, <NUM>, <NUM>, <NUM>) for exerting a compression force on the container (<NUM>); and
an actuator control (<NUM>) for varying the compression force.