Abstract:
A device for cardiocirculatory assistance, also named a ventricular assist device, includes a haematic pump ( 50 ) with a pump body ( 7 ) having an inner space ( 21 ) defined by a rigid structure ( 51 ) and a pair of mobile membranes ( 16, 17 ) alternately driven in opposite directions by alternately positive pressure and negative pressure gas supplied to recesses ( 19, 20 ) surrounding the two membranes by means of an external pneumatic force generating unit ( 1 ). This device achieves excellent operation results while maintaining a reduced size and reduced weight.

Description:
FIELD OF THE INVENTION 
   The present invention relates to a device for cardiocirculatory assistance which is obtained by means of a special haematic pump driven by pneumatic energy. 
   BACKGROUND OF THE INVENTION 
   It is known that in heart-surgery practice some devices for the mechanical assistance of the heart are used, commonly indicated as “ventricular assist devices (V.A.D.)” or “total artificial hearts (T.A.H.)”. 
   These are essentially devices adapted to mechanically pump blood to produce pulsing or continuous blood flows. 
   Such devices are employed to solve acute reversible heart failure clinical cases (as infarct or myocarditis) or are used to support the circulatory function while waiting for a heart transplant. 
   There have been for years several both left and right ventricular assist devices or biventricular assist devices (T.A.H); some of them are commercially available, others have been developed only at an experimental stage. 
   In most cases, the currently existing devices exhibit some difficulties to be accommodated inside the chest, because of problems both with size and weight. 
   Other drawbacks that can be found therein are caused by their inner configurations and their pumping modes, which may induce haemolysis or blood clot formation. 
   Some solutions currently use a balloon (ventricular sac) or a single membrane as a blood contacting flexible element. These devices, as their inherent feature, do not allow blood circulation in the pump, thus avoiding the above mentioned problems. Another negative aspect of these currently available solutions is represented by the weight and size of the driving units associated with the pumping device, which render the portability, namely the easy transport by the patient, essentially non-existent. 
   SUMMARY OF THE INVENTION 
   In view of this state of the art, it was the object of the present invention to obtain a device for cardiocirculatory assistance exhibiting high reliability and having reduced weight and size. 
   In accordance with the present invention, such an object is achieved by a device for cardiocirculatory assistance characterised in that it comprises a haematic pump with a pump body having an inner space defined by a rigid structure and a pair of opposed position mobile membranes, a pair of rigid lids defining around the mobile membranes respective recesses supplied with alternately positive pressure and negative pressure gas by an external pulsing pneumatic energy generating unit, so as to alternately draw the mobile membranes reciprocally together and apart in order to respectively decrease and increase the volume of the inner space of the pump body. The devise also comprises a blood input tubing in the inner space of the pump body, a blood output tubing from the inner space and a pair of spontaneously opening one-way valves respectively associated to the tubings in order to allow, in combination with the volume variation of the inner space of the pump body induced by the alternate motion of the two membranes, a continuous filling and evacuating cycle of the inner space of the pump body, therefore generating a blood flow rate depending on the number of pulses provided by the external pneumatic energy generating unit. 
   By means of its special two-membrane pumping system, the haematic pump comprised in the device according to the invention, allows to obtain a particularly efficient and effective haematic flow, as well as exhibiting high reliability, also obtaining on the whole a device having no inertial sussultatory motions, where the periodically pulsing haematic flow provides pulses, which may be synchronous or asynchronous with respect to the heart pulse, the whole driven by an external pneumatic energy generating unit which may have reduced weight and reduced size. 

   
     BRIEF DESCRIPTION OF THE DRAWINGS 
     A specific embodiment of the invention will now be described in detail, by way of a non-limiting example, with reference to the attached figures, in which: 
       FIG. 1  shows a diagrammatic view of a device for cardiocirculatory assistance according to the invention; 
       FIG. 2  shows a side view of the haematic pump comprised in the device of  FIG. 1 , with the membranes shown in a resting position; 
       FIG. 3  shows a longitudinal section of the haematic pump in a plane parallel to that of  FIG. 2 , with the pump shown in the resting position; 
       FIG. 4  shows a similar longitudinal section of the haematic pump during a delivery or compression step; 
       FIG. 5  shows a similar longitudinal section of the haematic pump during a suction or filling step; 
       FIG. 6  shows the haematic pump in longitudinal section according to line VI-VI of  FIG. 3 ; 
       FIG. 7  shows the haematic pump sectioned along line VII-VII of FIG.  3 . 
   

   DETAILED DESCRIPTION OF PREFERRED EMBODIMENTS 
     FIG. 1  shows a diagrammatic drawing of the various components of a device for cardiocirculatory assistance according to the present invention. 
   The device comprises a pneumatic energy (gas) generating unit  1  supplied by a battery set  2 , which transmits pneumatic energy to a tube  5  positioned inside the body of the patient (generally indicated by reference number  100 ) through a flexible tube  3  and a transcutaneous implant  4 . 
   The inner tube  5  is connected to an input  6  of a haematic pump  50  having a pump body  7  in a haemocompatible material sealed between two rigid protective lids  8  and  15 , obtainable in a material compatible with the body tissues of the patient, reciprocally fixed so as to form a single piece, as shown in  FIGS. 1 ,  2  and  3 . 
   Between the two lids  8  and  15 , the pump body  7  envisages two flexible inner membranes  16  and  17  ( FIG. 3 ), defining a substantially cylindrical shaped inner space  21  together with a remaining rigid structure  51  ( FIGS. 3 ,  6  and  7 ), communicating with an inlet  9  and an outlet  10 , to which respective haemocompatible tubings  13  and  11  commonly used in heart surgery techniques are connected ( FIG. 1 ), where there are housed two automatic one-way valves  12  and  14 , named spontaneously opening valves, commonly used in valve replacement surgery. 
   Under the control of the automatic valve  14 , tubing  13  withdraws blood from an atrium of the heart (e.g. through an atrioventricular tube) or from other areas of the cardiovascular system and inputs it in the inner space  21  of the haematic pump  50 , which pushes the blood itself to the blood vessels of the systemic circulation and pulmonary circulation through the other tubing  11  provided with automatic valve  12 . 
     FIG. 3  shows a longitudinal section of the haematic pump  50  in a resting position. 
   In operation, through connection  6  and an inner channelling  18  of the pump body  7 , a gas from generator  1  and from tubes  3  and  5  (named operating gas) is initially inputted inside two recesses  19  and  20  defined between the two flexible membranes  16  and  17  and the two rigid lids  8  and  15  and then in-taken from the recesses. 
   The blood in inner space  21  of the pump body  7  does not contact the compressed gas generated by the pneumatic energy external unit  1  as it is separated from it by means of two mobile membranes  16  and  17 . 
   In turn, the two lids  8  and  15  avoid contact between the working gas and the body tissues of the patient. Their resistance also allows an efficient protection against knocks and any handling action exerted on pump  50 . 
   As shown in  FIG. 4 , haematic pump  50  comprises a compression or delivery step, in which gas deriving from the pneumatic energy external generator  1  of  FIG. 1  is pushed inside the two recesses  19  and  20  through channelling  18 . The compressed gas inputted in recesses  19  and  20  pushes the mobile membranes  16  and  17  towards each other and towards the inside of the pump thus inducing a decrease of the inner space  21 . 
   The blood contained therein no longer finding a space to occupy, is thus forced to leave the pump body through the only available aperture, the one obtained at outlet  10  ( FIG. 6 ), from which tubing  11  provided with automatic one-way valve  12  ( FIG. 1 ) allows the blood to flow to the systemic or pulmonary circulation. 
     FIG. 5  shows a section of the haematic pump in a following suction or filling step. In this step, gas is in-taken from the two chambers  19  and  20  by the external pneumatic energy generator  1  through conduit  18  and tubings  5  and  3  connected thereto by means of connection  6 . 
   The evacuation of the working gas from the two chambers  19  and  20  determines the motion of the two membranes  16  and  17  in a direction opposite to that of the compression or delivery step, namely an outward direction, thus inducing an increase of the inner space  21  of the pump body  7 . 
   This exclusively blood-occupied volume draws from the outside, by increasing its capacity, more blood that enters space  21  through inlet aperture  9  and tubing  13 , where valve  14  of  FIG. 1  automatically opens to allow a one-way blood flow directed inside the pump body. 
   As already mentioned, both valves  14  and  12  are one-way flow and spontaneously opening, namely their aperture occurs automatically depending on the blood flow direction in the respective connection tubings  13  and  11 , and therefore on the pressure in inner space  21  of the pump. More precisely, valve  14  opens if the pump draws blood into inner space  21  and closes if the blood flow tends to invert its direction, whereas valve  12  operates in an opposite way, by closing when blood moves inside the pump and opening if blood is pushed out of the pump itself. 
   The above described operation steps with reference to  FIGS. 5 and 4  alternately follow each other so as to generate a continuous blood suction and delivery cycle, which allows the haematic pump to operate as a mechanical servo system to the heart. 
   During such cycle, the two membranes  16  and  17  are subjected to homogenous radial tensions, which render the cycle itself regular and preserve the membranes from premature wear. 
   Note that the substantially cylindrical inner conformation, with circular or elliptical bases, of space  21 , clearly shown in  FIG. 6 , in combination with an appropriate processing of its inner surface, allows blood to move from the inlet  9  to the outlet  10  following a circular path skimming the inner wall of the space itself, exerting a continuous washing of the same and avoiding the formation of stasis and particular stresses for the molecules and substances residing in the blood, which may generate dangerous circumstances for the health of the patient. 
   Furthermore, the two mobile membranes  16  and  17  are designed so that the value of their inertial masses and accelerations is such that the inertial forces existing in the haematic pump cancel each other so as to generate a sum of forces equal to zero or very close to zero, thus avoiding sussultatory motions caused by inertial forces. 
   The function of the operating gas and the role carried out by external pneumatic energy generating unit  1  of  FIG. 1  are now apparent. The operating gas used to compress inner membranes  16  and  17  of pump  50  is shifted by periodic alternate motion in alternate input/output direction to/from the haematic pump by means of a positive pressure/negative pressure generated by external unit  1 . 
   This external unit is obtained with reduced weight and reduced size features so as to facilitate its transport by the patient easily without using special equipment such as trolleys or other voluminous equipments. It may be carried on the shoulder back or attached to the trouser belt of the patient by appropriate systems. 
   In brief, the device for cardiocirculatory assistance according to the present invention is a device which is capable of shifting sufficient blood amounts, producing a pulsing blood flow with variable flow rate depending on the number of pulses per minute generated by the external pneumatic energy unit by means of the operating gas. 
   The special inner configuration of the pump allows a particularly efficient and effective haematic flow also to be obtained as well as on the whole a device having no inertial sussultatory motions, where the periodic pulsing haematic flow provides pulses which may be synchronous or asynchronous with respect to the heart beat, the whole being driven by a small sized pneumatic energy generator. 
   The pulses generated may thus create the blood pumping state having pulsing motion synchronous to the heart beat, which is a specific feature of this invention. 
   It is known that the pumping of blood having a pulsing motion is much more physiological than the continuous motion obtained by currently existing centrifugal pumping systems. 
   Many minor modifications, not affecting the general principle at the basis of the present invention may of course be made to the above described device by way of example. 
   Specifically, valves  12  and  14  of  FIG. 1  may be housed inside blood suction and delivery inlet  9  and outlet  10 , instead of being inserted in tubings  11  and  13 . 
   The inlet  9  and outlet  10  may in turn be reciprocally angled or parallel. 
   When the two membranes are in resting position, they may lie on reciprocally parallel planes, as shown in  FIG. 3 , or on reciprocally diverging planes in the direction of inlet  9  and outlet  10 .