The present invention relates to a peristaltic pump of the type comprising at least two elastically compressible hose conduits, each having an inlet and an outlet end and mutually identical cross section; a support means for each 5 ' hose conduit, said support means having a curved surface which is similar for all the hose conduits and which is engaged by the associated hose conduit with at least a por¬ tion of the length thereof; at least two synchronously drive rotor units, one for each hose conduit, which are provided 10 with the same number, at least two, of symmetrically arrange and- mutually at the same distance from the rotor axis opera- ring pressure means for local compression of the hose con¬ duit corresponding to each rotor unit against the curved surface of the associated support means, whereby the number 15 of pressure means on each rotor unit and the extension of said curved surface on the support means of each hose con¬ duit, seen in the longitudinal direction thereof, are mutual ly so adapted that continuously one of the pressure means on each rotor unit compresses the corresponding hose conduit fo 20 preventing fluid flow through the hose conduit in a direc¬ tion opposite to the direction of rotation of the rotor unit.

Peristaltic pumps of this type generates a pulsatile fl in which the pulsation frequency is related to the number of revolutions of the pump. Within certain areas however, a 25 pulsatile flow is utterly disadvantageous. Thus, peristaltic pumps with pulsatile flow have blood-damaging properties

I

(hemolysis and trombocyte aggregation) when they are used fo blood pumping in heart-lung machines and dialysis equipment. Neither are peristaltic pumps with pulsatile flow suitable » 30 when certain materials e.g. putty material, are to be uni- formly distributed over certain surfaces, e.g. floors.

The object of the present invention is to eliminate said drawbacks and provide a peristaltic pump permitting a non- pulsatile flow. This as arrived at according to the inven- 35 tion by that at least the outlet ends of the hose conduits forming part of the pump are joined into a common outlet in a manner known per se; and that the synchronously driven

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rotor units are angularly displaced relative to each other such that their pressure means operate with phase displace¬ ment on each hose conduit, whereby the flow pulsations gene¬ rated in each hose conduit will at least to a substantial ex- tent neutralize each other at the common outlet of the hose conduits.

The peristaltic pump according to the invention genera¬ tes a non-pulsatile flow at constant number of revolutions and is relatively blood compatible. Therefore, the pump is especially suited for control by means of a feed-back circuit and also to be utilized for pumping blood in experimental re¬ search on animals, in which requirements for accurate flows and pressure levels and pressure variation configurations should be obtained in combination with low hemolysis. Since the pump through its construction is relatively blood compa¬ tible, it is also applicable as a pump in heart-lung machines and dialysis equipment but may find many other applications in laboratories and in the industry.

In each of the conduit branches connected in parallel an approximately sinusoidal flow is generated, said flows being generated symmetrically displaced relative each other and superposed on each other they generate a non-pulsatile flow. By letting the pump be controlled by e.g. a normal hear -pressure curve, the pressure curve thus produced by the pump may be brought to almost completely imitate a normal heart- -pressure curve. Thus, the pump permits normal pulsatile as well as constant pressure perfusion and non-pulsatile constan -flow perfusion at the desired flow and pressure levels respe tively. Also, other pressure configurations such as sinusoidal triangle, impulse, etc. may be obtained by controlling the pump by means of a functional generator with such curve vol¬ tage variations. This pump is- through its features suitable for use in the most experimental investigations on animals within peripheral circulation physiology, but also suitable as an infusion pump with constant flew velocity also for very small flows, which is made possible by selecting small dia¬ meters in the conduit branches. Additionally, the pump is, as mentioned, blood compatible by its special design as descr bed below.

The invention will be further described below with refe¬ rence to the accompanying drav/ings, in which fig. 1 is a perspective view of a peristaltic pump accor ding to the invention and here with three parallel rotor unit 5 fig. 2 is a section through fig. 1 perpendicular to the rotor axis; fig. 3 and 4 is a plan and side view respectively, of a conduit system running through the peristaltic pump; fig. 5 is a principal sketch of the pump and its control 10 units for providing constant pressure or constant flow per¬ fusion; fig. 6 shows original registrations of pump-generated . pressure curves; fig. 7 is a perspective view of members in a peristaltic 15 pump according to the invention with two parallel rotor units fig. 8 is a section through an alternative embodiment of the pump according to the invention; and fig. 9 is a section along the line IX-IX in fig. 8. The principal construction of a peristaltic pump means 20 that medium is led through a flexible tubing without bringing any other part of the pump in contact with the medium. A pres sure means occludes the tubing against a curved surface on a support means and the pressure means forces the medium ahead during rotation. The flow thus produced is pulsatile and the 25 average speed thereof substantially proportional to the numbe of revolutions of the rotor.

The principal construction of a peristaltic pump accor¬ ding to the invention is shown in fig. 1. This pump 24 is pro vided with three parallel rotor units 1, of which each has 30 three pressure means 3 which are ¹ arranged symmetrically rela- * tive to each other and relative to the rotor axis 2 and which . during rotation of the rotor move along a curved surface 4 on the support means and by deformation of a medium-carrying con duit system 5 displace the medium therein (see also fig. 2) . 35 The pressure means 3 comprise rotatably mounted rollers and are in each rotor unit distributed symmetrically around the rotor axis 2, i.e. while there are three pressure means 3 in each rotor unit 1, the angle between each pair of pressure means is 120 . The pressure means 3 in one rotor unit 1 are

also displaced relative to the pressure means 3 in another rotor unit 1 such that the totally nine pressure means 3 are positioned symmetrically at an angle of 40 between each pres sure means. The conduit system 5 which e.g. by means of screw joints 28 is mounted on the support means, comprises in the embodi¬ ment shown three parallel conduit branches 6 of which each and everyone is associated with a rotor unit 1. The conduit branches 6 are also joined into a common outlet 7. Medium is fed to the outlet while the pressure means 3 define medium- -containing spaces 8 in the conduit branches 6, displace said spaces towards the outlet 7 during successive increase of the volume thereof, open said spaces successively towards the out let and close said spaces towards the inlet 9 of the conduit branches.

The conduit branches 6 of the ^" conduit system 5 are made of elastically compressible hose conduits of similar cross section and each conduit branch 6 engages with at least a portion of its length a curved surface 4 which is the same for all the conduit branches. The rotor units 1 are driven synchronously and their symmetrically arranged and mutually at the same distance from the rotor axis 2 operating pressure means 3 permit local compression of the conduit branch corres ponding to each rotor unit 1 against the associated curved surface 4, whereby the number of pressure means 3 on each rot and the extension of said curved surface 4 for each conduit branch 6, seen in the longitudinal direction thereof, are mutually so adapted to each other that always one of the pres sure means 3 on each rotor unit 1 compresses the correspondin conduit branch 6 for preventing fluid flow through the condui system 5 in a direction opposite to the direction of rotation of the rotor unit 1.

While at least the outlet ends of the conduit branches 6 ^"• re joined into the common outlet 7 and the synchronously dri ven rotor units 1 are angularly displaced relative to each other as stated above such that their pressure means 3 operat with phase displacement on each conduit branch, the flow pul¬ sations generated in each conduit branch will at least to a substantial extent neutralize each other at the common outl<

of the conduit branches.

The successive increase of the volume of the medium-con¬ taining spaces 8 and their gradual opening towards the outlet 7 and closure towards the inlet 9 is provided by giving the curved surface 4 a special design. Thus, a first part 10 of ^■ the curved surface 4 may be circular in shape, run at a cons¬ tant distance from the rotor axis 2 and have a length corres¬ ponding to or somewhat exceeding the distance between two adja cent pressure means. Thereafter, the circular part 10 may (towards the outlet 7) be transformed into a rear part 11, the distance of which from the rotor axis 2 successively increases, towards the outlet 7. While the curved surface ini¬ tially very slowly but successively increases the distance from the rotor axis 2, but at an ever increasing rate with decreasing distance to the outlet it will successively and initially to a very small extent per degree of movement to¬ wards the outlet open said spaces 8 faster in the end until they are completely open when the pressure means 3 leaves the curve _^ surface 4. The rear part 11 of the curved surface 4 is preferably made elliptical, but may also have another shape. In order to prevent backward flow at any time, in the embodiment shown, at least the first 120° of the curved sur¬ face 4 defines the circular fore part 10, in which part total compression, i.e. occlusion of the conduit branches is attai- ned. Thus, the gradual release of the pressure means 3 from the curved surface 4 in the elliptical part 11 generates flow oscillations from each rotor unit 1 which are substantially sinusoidal. The arc of the curved surface 4 preferably comp¬ rises totally about 180° (see fig. 2 ) , but of course a longer arc is also possible.

At the outlet 7 each conduit branch 6 generates a substan tially sinusoidal flow, q, , q_ and σ.. respectively, with a mutual phase displacement θ of 360°/3, i.e. 120°. Provided the flow is completely sinusoidal with the amplitude q and average flow q from each rotor unit 1, q, , q ₂ and q ₃ may be expressed as:

^q l ^{= q} m ^{+ g} _P P ^{sin wt} ' ^q 2 ^{= q} rπ ^{+ q} p ^{sin (wt +} ^ ^{] ; q} 3 ^{= q} m ^{+ q} p

and the total flow (Q) as:

Q = <ι_ + g ₂ ^{+ q} 3' giving

Q = 3q + q sin wt + sin (wt + θ ) + sin (wt + 20 ) = 3q _m + q Vl + 4 (cos θ + cos Q ) - sin (wt + 0 ) .

Insertion of = 120° gives

i.e. a non-pulsatile flow. p, , „ and p_ represent the pressure in each conduit branch 6 and p. the pressure in the outlet 7 . R, , R- and R_ represent the flow resistance in each conduit branch 6 from the forward, totally occluding pressure means 3 to the com¬ mon outlet 7. R. represents the flow resistance at the outle 7 distally of the connecting point of the three parallel con duit branches, including the prefounded vascular bed. . can be expressed as:

— ^P l± + _ ^P _2±. + _ ^P _-3±

P _ά = ^R l ^R 2 ^R 3

1 1 1 i

} ^m •-_. " _m /

The conduit branches 6 are designed such that R, , R_ an R are small compared to R. , wherefore p, can be simplified:

+

^ "ΪL

With a mean pressure p_ and a pulse pressure amplitude

171 p for all three rotor units, Ό, , D_ and p., can be exDressed p '1 -2 ^3 as:

^P l ^{= p} m ^{+ p} p ^{Ξin wt? p} 2 ^{= p} m ^{÷ p} p ^{sin (wt ÷} ® ^{) ; p} 3 ^{= p} m ^{+ p} and for0 = 120° *± τr. - S ^E i.e. a non-pulsatile pressure. ( OMPI

In order to be able to control the pump-produced pres¬ sure or flow via a negative feed-back circuit and create exac pulse configurations, the pump must produce a uniform and virtually non-pulsatile flow at a constant rotor speed. Pro- vided the flow oscillations from each rotor unit 1 are sinuso dal, which has been accomplished with the present pump, it follows from the mathematical calculations above that the pre sent pump will deliver a flow and a pressure which are substa tially constant at the outlet 7 at a constant rotor speed. This design thus fulfils the criteria for allowing servo-

-control of the pump via a negative feed-back circuit permit¬ ting the pump not only to provide a constant mean pressure or constant flow perfusion within a wide pressure or flow range, but also all types of pressure and flow configurations. It is a well-known fact that pump perfusion of blood may damage the blood cells. The break-down of red blood cells, designated as hemolysis, but also the degree of trombocyte destruction, are usually taken as a measure of such blood damage. The major factors causing blood damage when using pumps of so called roller type have been shown to be related to the conduit branch material, the smoothness of the inner surfaces of the conduit branches, too high blood flow velo¬ cities, too small conduit branch diameters, too high rotor speed and frequency of conduit branch occlusions. Blood dama- ging influence of these factors has been minimized in the pre sent pump device 24 by its special design as described below. The conduit branches 6 are made of a blood compatible, elasti material, e.g. silicone rubber, with an inner surface covered by an even more tissue compatible material. The conduit branc are moulded in one piece with the inlet 9, which is divided at the end of the curved surface, and are drained at the com¬ mon outlet 7. The outlet 7 also comprises a suitable device 1 for registration of the pressure of the medium in the outlet 7 (see fig. 2) . The device 12, which comprises a pressure met 25, consists of a side outlet 26 to the outlet 7 with a small diameter in order to as little as possible interfere with the fluid. All conduit branches 6 brought in contact with blood are moulded in such a way that extremely smooth inner surface are obtained. Flow velocity and the frequency of conduit bran

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occlusions have been minimized by choosing a relatively large rotor diameter (70 mm in this embodiment) and a large inner diameter for the conduit branches (6 mm) , whereby a very low rotor speed is reached. With this embodiment the pump gives a very low hemolysis at flows up to about 40 ml/min., more exactly defined < 0,008 g/1 of he oglobine added to plasma at each passage of the blood through the pump, which is to be compared with a value of > 0,14 g/1 with a conventional blood perfusion pump (Harvard Variable Speed Peristaltic Pump, Mode 1210) at the one and same given test.

The physiological criteria mentioned above, are fulfille by the present pump for blood flows up to about 40 ml/min. Maximum flow however, is about 300 ml/min. , but for these flo the degree of blood damage is larger. Fig. 5 is a block diagram of the pump 24 and its control units for constant pressure or constant flow perfusion. Cons¬ tant flow perfusion is accomplished by setting a switch 13 in position 1 yielding a summation of a signal proportional to the rotor speed and a signal representing the desired blood flow. The resulting signal is fed into a PID-regulator 14.

The rotor speed signal is obtained from a tacho-generator 15 placed on the drive shaft of the rotor. The sum of a signal proportional to the pump produced pressure (derived from a se rate pressure transducer 16) and a signal representing the de sired arterial pressure (referred to as the reference signal) then gives an input signal via the PID-regulator 14 and a pow amplifier 17 to the motor unit 16 comprising a gear box 19 an producing a rotor speed which at any moment will keep the pre sure at the set constant level. Thus, the desired blood flow or blood pressure level can be obtained by manual variation o a reference signal. By separating the reference signal into a DC and an AC component, a constant average pressure or consta flow (DC-component) of any desired magnitude can be produced independently of a superimposed pulse pressure (AC-component) With an appropriate selection of motor and gear ratio, this pump is capable of reproducing with great accuracy the normal arterial pulse pressure curve both during the upstroke and the pressure decrease. The pressure decrease during the diastolic phase of the πulse pressure curve can, however, no

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be obtained simply by an immediate stop of the rotor. For thi purpose the rotor direction has to be reversed for a very short period of time. Such transient backwards rotation is obtained with the aid of a special electronic unit 20 which consists of a highpass-filter in combination with a recti¬ fier.

With the aforementioned construction, the pump 24 is capable of reproducing the normal arterial pressure curve up to a frequency of about 4 Hz by using the normally undamped cardiac pulsations (registered via a separate pressure trans¬ ducer from a catheter in a systemic artery) as the AC-compo- nent in the reference signal. Other types of AC-signals can alternatively be applied, e.g. sinusoidal, triangle, impul¬ ses, step functions etc., which preferably are obtained from a function generator.

Fig. 6 illustrates some examples of how the pump 24 described above has been used to reproduce various configura¬ tions of pressure curves. Panel A, the upper curve, shows the normal undamped systemic arterial pressure registered. from a cat's brachial artery with a Statham pressure trans¬ ducer. The lower curve in Panel A illustrates the simulated curve produced by the perfusion pump device. Note the close resemblance between these two curves both for slow respira¬ tory variations and the cardiac induced pulse pressure varia- tions. Panel B shows how the pump can be used to produce a sinusoidal pressure curve, Panel C a positive and a negative pressure impulse and Panel D a positive and a negative pres¬ sure step function.

The present invention is not limited to the embodiment described above. Thus, the pump described is made for blood perfusion for experimental research on animals, but the pump may be used for all types of perfusion, not only blood per¬ fusion in heart-lung machines and dialysis apparatuses, but also for perfusion of other flows than blood within all flow areas. A modification of the above pump has been constructed for constant flows as small as 1 • 10 ml/min. The pump may comprise two or more than three rotcr units 1 and each rotor unit may have two or more than three pressure means 3. Furthe more, the conduit system 5 may of course consist of two rro^j^_όts.^_

more than three conduit branches 6. The length of the curved surface 4 may vary and so may also the circular and ellipti¬ cal parts 10, 11 respectively. In order to be able to cont¬ rol the pressure of the pressure means 3 on the conduit branc 6 and see to that each pressure means presses against the co duit branches with the same force at any given point along th curved surface 4, the rotor units 1 are by means of setting devices 22 and 23 of suitable construction also adjustable laterally, vertically and in various inclined positions rela- tive to the curved surface 4. The vertical and lateral adjus ment is carried out e.g. with screw devices 22 displaceable in long holes 21 and the inclined position is set preferably by means of screws 23.

The pump illustrated in fig. 8 and 9, is intended for pumping floor masses. It has two rotor units 1 which coope¬ rate with one conduit branch 6 respectively and which each have two pressure means 3 in the form of radially projecting cam means. The angle therebetween is 180 and the angle betw one cam means in one rotor unit and the cam means closest thereto in the other rotor unit is 90°. The function of the pump corresponds with the pump illustrated in fig. 1 and 2.