WO1996018425A1

WO1996018425A1 - Injector

Info

Publication number: WO1996018425A1
Application number: PCT/GB1995/002969
Authority: WO
Inventors: William Anthony Denne
Original assignee: Bailey, William, John; Reid, Dominic, Augustine
Priority date: 1994-12-17
Filing date: 1995-12-15
Publication date: 1996-06-20
Also published as: AU4268096A

Abstract

A preferred example of an injector (1) for needle-less administration of a drug has a rigid member (43) and a conical shell piezoceramic actuator (41) which is anchored to the member (43) around the skirt of the cone. The injector (1) has a pump chamber (54) having an inlet through which a liquid drug can enter and an outlet with a non-return valve (55) through which the liquid drug can be ejected. Operation of the actuator (41) reduces the volume of the pump chamber (54) thereby ejecting a droplet of liquid drug out through the outlet at sufficient speed to pierce a recipient's skin.

Description

INJECTOR

The present invention relates to an injector.

The subcutaneous administration of drugs has been practised for several centuries. The advent of the hypodermic syringe greatly expedited drug injection and the method is currently in very widespread use.

In the last 50 years, liquid drugs have been administered by jet injector. With this technique, a quantity of liquid drug is fired at the recipient with such velocity that the liquid jet penetrates the skin. This has the advantage of repeated dosing for mass inoculation without any need to change or sterilise the needle between shots. There is also the potential for pain-free injection. The currently available devices accelerate the drug by means of a spring action, compressed gas or explosive devices. Both the spring and gas powered injectors deliver the dose with such speed that considerable subcutaneous soft tissue damage occurs. The resultant bruising is frequently more painful and results in a longer lasting pain than a hypodermic needle injection. In addition, the damage causes scar tissue which is of considerable significance to repetitive users such as diabetics. For less frequent recipients of injections, fear of the conventional hypodermic needle is not insignificant. However, the obvious effort in cocking a spring driven device and the resounding clunk on firing are as disconcerting to the recipient as a needle prick. Similarly, gas driven devices tend to hiss alarmingly.

In summary, there is a significant need for a silent jet injector that can introduce the dose as slowly as a hypodermic needle but without the obvious pin prick of the needle. According to the present invention there is provided an injector for needle-less administration of a drug, the injector comprising: a rigid member; an actuator; a pump chamber between the rigid member and the actuator, the pump chamber having an inlet through which a fluid can enter and an outlet through which the fluid can be ejected; and, a non-return outlet valve at the outlet, whereby operation of the actuator reduces the volume of the pump chamber thereby ejecting fluid in the pump chamber out through the outlet and non-return outlet valve at a speed sufficient to pierce the skin of a recipient.

The pump may eject a measured dose of flowable drug at high enough velocity to penetrate the skin. The pump may be repetitive in action and may cycle at ultrasonic frequency to reduce noise. The pump may be a reciprocating pump such that the drug may be ejected as a series of high velocity droplets. The actuator may be piezoelectric, electrostrictive, magnetostrictive or other means.

In one preferred example, the actuator is a thin concave shell of PZT ceramic, the apex or upper third of which carries electrodes so that excitation causes the shell to contract, the bottom portion or lower third of the shell being rigidly bonded with epoxy resin to the rigid member which fits intimately to the bottom portion of the shell but which diverges from the shell under the electrode area to form the pump volume. The pump volume is preferably filled with resilient material except for a fine axial tube which serves as the pump chamber. As the shell contracts on excitation, the resilient material is displaced radially inwards to displace the walls of the pump chamber and thereby reduce its volume. The modulus of the resilient material and the profile of the pump chamber should be chosen to permit the resilient material to move freely rather than be constrained by cumulative radial stress. A conical shape offers a number of specific advantages which will be described more fully below. The conical actuator may be rigid]y onded about its periphery to the rigid member such that there is a gap which forms the pump chamber between the inner surface of the actuator and the rigid member in the apical region.

The pump chamber may be formed by a defor able elastomeric moulding that incorporates a high speed valve or valves. Preferably this is a silicone rubber moulding incorporating non-return valves. A valve orifice may be in the form of a slit or system of slits to generate drops with knife-like leading edges. An embodiment comprises a modified piston and cylinder arrangement. The pump volume is preferably substantially filled with deformable silicone rubber in such a manner as to effectively eliminate most of the dead space and make the pump self-priming. Conveniently, there is a relatively long rigid small diameter axial feed leading to the pump chamber and a high speed non-return valve adjacent to the said chamber. A rapid compression stroke generates a high cylinder pressure which ejects fluid through the non-return valve rather than down any feed tube due to the difference in fluid inertia. A high pressure pulse propagates down the feed tube, but with appropriate valve flexibility, or appropriate electrical drive waveform such as sawtooth, there is no subsequent low pressure that might cause cavitation in the drug and essentially no interruption of feed flow. The device may be driven below mechanical resonance by a regenerative circuit so that the drive is electrically efficient yet the volume of drug dispensed is not dependent on the quality factor of a mechanical resonance and the dispensing rate may be decided arbitrarily.

The provision of a silent high pressure micropump is not straightforward. The existing means are far from silent. Electric motors, solenoids and the like are also noisy. Ultrasonic devices may by definition be extremely quiet. Piezoelectric ceramics may produce high pressures, are extremely compact and may be efficient electromechani al actuators if correctly driven. They are significantly more effective than electrostrictive or magnetostrictive elements. Operation at very high frequency is no problem and the use of silicone mouldings provides the possibility of extremely high speed hydraulic valves. However, the piezoelectric devices are reciprocating in nature, the displacements are extremely small and the devices are very rigid.

The reciprocating nature of the piezoelectric actuators is in fact an advantage in the context of the present invention. Rather than ejecting the drug dose in a single slug of high velocity fluid, the drug may be dispensed as a series of high velocity droplets. This provides two major advantages.

The first advantage is that the dose rate may be reduced without reducing the droplet velocity or requiring prohibitively fine nozzles. In this manner, the drug may be injected at a speed comparable with a needle injection so reducing soft tissue damage.

The second advantage accrues from the fluid dynamics of fast droplet impact. When a jet impinges on a surface, it rapidly establishes an equilibrium flow such that the interface pressure is half the liquid density times the jet velocity squared. This pressure must significantly exceed the maximum skin strength to ensure penetration of the jet rather than formation of a stable high velocity surface flow. A single droplet impinging on a surface will have a similar effect at low velocity, but at higher velocity, the impact will produce a shock wave in the droplet which will travel at the speed of sound in the liquid. It may be shown that the maximum impact pressure becomes half the density of the liquid times the droplet velocity times the speed of sound in the liquid. Much higher impact pressures may therefore be generated by discrete drops than a continuous jet. A velocity of 100 m/s may be used instead of 200 m/s required for a continuous jet injector. This results in a quarter of the power input, which is significant for battery powered devices. The greater surface penetration is associated with reduced range in the underlying tissue. The latter arises not just because of the lower velocity but also because the drag on a spherical droplet is much higher than on a cylindrical column of fluid. This is important because many injection sites rely on limited but well controlled penetration depth.

The rigidity of the piezoceramic is also an advantage.

To produce a droplet velocity of 100 m/s requires a driving pressure of 50 bar. This stress corresponds to 75 icrostrain in PZT which is well within the maximum working tensile limit of around 300 microstrain.

There are a number of ways of driving the PZT. The applied electric field may be parallel to the direction of actuation, or d₃₃ mode. It may be perpendicular to the direction of actuation, or d₃₁ mode. There is even a shear displacement mode, d₁₅. The actuator may be used with a mechanical movement magnification device, such as a horn or unimorph arrangement. The magnification devices limit the frequency-displacement product and so reduce pumping rate. The shear mode transducer is not straightforward to use.

This leaves d₃₃ and d₃₁ actuation.

With d₃₃ operation very large voltages across a single plate are required to produce significant movement. Typically, 2 kV will produce 1 micrometre displacement. For practical operation it is desirable to use a stack of many thin plates at lower voltage and such devices are available as monolithic multilayer actuators. Typically less than 50 V is required to produce 1 micrometre displacement. With d₃₁ operation, the plate can be very thin so that a modest electric field produces a large strain. A long plate will then produce a large displacement. The disadvantage of this arrangement is that the frequency- displacement product is much lower than the d₃₃ case. The outlet may be a slit or system of slits so that rapid opening may occur without imparting significant kinetic- energy to the valve lips. The slit opening may impart a knife-like leading edge to the drop to enhance skin penetration.

Examples of the present invention will now be described with reference to the accompanying drawings, in which:

Fig. 1 is a front elevation of a first example of an injector of the invention;

Fig. 2 is a cross-sectional view on A-A of Fig. 1;

Fig. 3 is a cross-sectional view on B-B of Fig. 1; Fig. 4 is a cross-sectional view on C-C of Fig. 3;

Fig. 5 is a detailed cross-sectional view of the pup chamber of the example of Fig. 1;

Fig. 6 is a cross-sectional view of a second example of an injector of the invention; Figs. 7a and 7b are detailed cross-sectional views of the injector of Fig. 6 showing operation of the injector;

Fig. 8 is a cross-sectional view of a third example of an injector of the invention;

Fig. 9 is a detailed cross-sectional view of the injector of Fig. 8 showing operation of the injector;

Fig. 10 is a perspective view of the tip of the injector of Fig. 8;

Fig. 11 is a diagrammatic representation of the contact between the injector of Fig. 8 and a patient's skin; and,

Fig. 12 is a detailed view of an actuator of a variation of the third example.

In Figs. 1 to 5, there is shown a first example of an injector according to the invention. The first example has a piezoelectric actuator operating in d₃₁ mode.

In the injector 1 of Figs. 1 to 5, a drug feed tube

2 is set in the base 3 of a rigid member 4. The rigid member 4 has a generally C-shape cross-section in which the base 3 and a top portion 5 project as can be seen in Figs. 2 and 3.

A blade-like piezoelectric ceramic actuator 6 is rigidly attached at one end to the top portion 5 of the member 4. The other end of the piezoelectric actuator 6 engages in a close-fitting slot 7 provided in the top surface of the base 3. The piezoelectric actuator 6 is sealed by a push fit into a silicone rubber moulding 8 provided in the slot 7. The interference between the piezoelectric actuator 6 with the sidewalls of the slot 7 is such that the rubber moulding 8 forms integral edge seals 9 around the portion of the piezoelectric actuator 6 that extends into the slot 7. These edge seals 9 may deform by shear under the influence of longitudinal movement of the piezoelectric actuator 6.

The silicone rubber moulding 8 is formed to provide a pump chamber 10 underneath the piezoelectric actuator 6 into which the feed tube 2 opens. The silicone rubber moulding 8 is formed to provide an integral inlet valve 11 of triangular cross-section between the feed tube 2 and the main part of the pump chamber 10. The silicone rubber moulding 8 is also formed to provide an outlet valve-cum- nozzle 12 of triangular cross-section at the end of the pump chamber opposite the feed tube 2.

The injector 1 may be fabricated by first producing the silicone moulding 8, which may also seal the feed tube 2 in position. The piezoelectric actuator 6 is then slid into position, pressed onto the silicone pump moulding 8 and bonded with for example an epoxy adhesive to the top portion 5 of the rigid member 4. As shown in Fig. 5, the triangular or V-shape cross-section of the moulded channel 10 forming the pump chamber 10 decreases in area with pressure applied to the actuator 6. It is desirable to ensure that the piezoelectric actuator 6 and the rigid member 4 have similar coefficients of thermal expansion, so that correct clearances are maintained, particularly at the pump region, at varying ambient temperature. Pressed alumina is a suitable material for the rigid member 4. In operation, the feed tube 2 is inserted into the septum of a drug ampoule (not shown) . A voltage is applied to the piezoelectric actuator 6 which extends to force air from the pump chamber 10 past the high speed non-return outlet valve 12. On removal of the voltage, the piezoelectric actuator 6 retracts, and air and then drug will be sucked through the feed tube 2 into the pump chamber 10 via the non-return valve 11. The injector 1 is thus self-priming. Once the pump chamber 10 is full of drug, a voltage can be applied again and the extension stroke of the piezoelectric actuator 6 will eject a high velocity drop of drug through the valve-cum-nozzle 12. Typically, a plate or blade 6 of PZT ceramic 0.2 mm thick will produce 100 microstrain in the longitudinal direction at 100 V. If the length of the blade 6 is 10 mm, there will be a displacement of the blade tip of 1 micrometre. If the blade 6 is 5 mm wide, the swept volume will be 0.001 mm . The longitudinal resonant frequency of the blade will be 75 kHz, which permits a volume dispensation rate of 0.075 ml/sec.

If the pump chamber 10 has a triangular or V-shape cross-section as shown in Fig. 5, it is possible to produce very small pump chamber cross-sections quite reproducibly, allowing the pump to be self-priming, whilst making it impossible to close the pump chamber 10 with initial loading from the piezoelectric ceramic plate 6.

A second example of the present invention is shown in Figs. 6 and 7. The injector 1 has a cup-shape base 21 which forms a cap to fit over a drug ampoule (not shown) and a base for a pump actuator 22. The actuator 22 is in the form of a cylindrical monolithic multilayer PZT actuator operable in d₃₃ mode. The actuator 22 and the base 21 each have an axial hole 23, 24 into which extends a sharpened tube 25, which may be a hypodermic needle. The needle 25 is cemented firmly in its seating in the base 21 with epoxy adhesive. In operation, the cup-shape base 21 fits over the drug ampoule and the needle 25 pierces the rubber septum of the ampoule to provide a fluid feed of drug. A cup-shape rigid member 26 fits over the actuator 22 and is bonded to it with a silicone rubber moulding 27. The inside face of the rigid member 26 is machined or otherwise formed so that there is a close fit around the periphery of the upper surface of the actuator 22 but significant clearance above it, the clearance increasing towards the centre. In the example shown, the inside of the face of the rigid member 26 is frustoconical. An axisymmetric hole 28 is provided in the centre of the rigid member 26. The rubber moulding 27 is moulded with a pointed pin in the feed tube 25 in order to provide a cavity 29 at the centre of the volume of the silicone moulding 27 between the profiled inner surface of the rigid member 26 and the upper surface of the actuator 22. The cavity 29 forms the pump chamber 29. Before extracting the moulded pin, it is forced through the part of the silicone moulding 27 formed on the upper surface of the rigid member 26 to form a high speed silicone non-return valve 30. The pin may have a polygonal cross-section which will produce for example a star-shape slit, which may have advantages in terms of rapid opening with low rubber momentum. The needle 25 extends through the PZT stack actuator 22 to the pump chamber 29 to provide a well defined capillary tube. The needle 25 is much more extensible than the PZT and will move freely with the actuator 22 when the actuator 22 is excited. Note that other shapes for the inner face of the rigid member 26 are possible, such as hemispherical.

The multilayer actuator 22 is typically 3 mm long providing a resonant frequency of 0.3 MHz. With sixty layers of 0.05 mm thickness, the actuator 22 will produce a 2 micrometre displacement at 70 V. The speed of sound in the captive silicone rubber 27 between the upper surface of the actuator 22 and the opposed inner surface of the rigid member 26 is approximately 1500 m/s. A 2 mm radius to the actuator 22 provides a 1.3 microsecond stress transit time in the rubber and a stroke volume of 0.025 mm . If the devjce is operated at 100 kHz, a dispensing rate of 2.5 ml/sec is attainable. Typically, operation in the d₃₃ provides much higher pumping rates than operation in the d₃₁ mode.

In operation, the actuator 22 distends axially on application of voltage as shown by arrows in Fig. 7a, thereby compressing the rubber 27 contained between the surfaces of the actuator 22 and the rigid member 26 against the inertia of the rigid member 26. The close fit of the rigid member 26 and the actuator 22 around the periphery of the pump chamber 29 may be engineered so that there is no significant extrusion of rubber from the chamber 29.

Instead, as shown in Fig. 7a, the rubber around the pump chamber 29 moves axially inwards. The pump chamber 29 collapses at the centre and air is driven from it through the nearby non-return valve 30. The voltage is then removed and the actuator 22 then retracts as shown in

Fig. 7b, thereby recharging the pump chamber 29 with air from the feed pipe 25 as the rubber around the pump chamber

29 relaxes as shown in Fig. 7b. After repeated application and removal of voltage, the air in the feed pipe 25 will be expelled, and drug is pumped through the non-return valve

30 in a similar fashion.

The high hydraulic pressure generated in the drug filled pump chamber 29 provides the drive to produce a high velocity droplet of the drug. The same pressure is resisted by the inertia of fluid in the feed pipe 25. The pressure propagates down the feed pipe 25 as an acoustic pressure pulse without significantly affecting the feed flow. If the valve 30 is sufficiently compliant, or a sawtooth electrical drive waveform is used, there is no subsequent negative pressure pulse which might otherwise cause cavitation in the drug. This deformation of the non¬ return valve 30 after drug ejection provides the low negative pressure required to maintain the continuous feed of drug.

At the high pressures and small displacements involved, t e bulk compressibility of the silicone rubber 27 may become significant. The volume 27 must be kept slim so that the small actuator displacement displaces the rubber rather than compressing it. The rubber deforms as a solid though it must act like a hydraulic fluid connection between the actuator 22 and the pump chamber 29. The rubber must have a sufficiently low modulus and the profile of the surface of the rigid member 26 opposed to the actuator 22 must be designed with care to ensure that there is no significant rigidity in the rubber which might otherwise cause degradation of pump performance by rubber compression.

A third example of an injector 1 is shown in Figs. 8 to 12. The injector 1 has a hollow shell piezoelectric ceramic actuator 41. In the example shown, the actuator is a hollow cone though other profiles are possible, such as hemispherical or ogival. Electrodes 42 are provided around the central (upper) portion of the conical actuator 41 and thus only the central portion of the actuator 41 is actuated by the electrodes 42, the balance of the skirt of the conical actuator 41 acting as an anchor. The actuator

41 mates with a rigid member 43, the mating surface 44 of which has two segments 44a, 44b. The shape of the outer segment 44a corresponds to the actuator 41. Thus, in the example shown, the outer segment 44a is conical and has the same angle of taper as the actuator 41 and is therefore a very good fit. The actuator 41 is bonded rigidly to the outer segment 44a with epoxy resin. The included angle of the inner or upper conical segment 44b is greater than that of the actuator 41, so there is an increasing clearance between the inner surface of the actuator 41 and the surface of the inner segment 44b with decreasing radius.

An axial hole 45 is provided in the rigid member 43 into which is bonded a sharpened tube 46, which may be a hypodermic needle. The tube 46 enters the septum 47 of a drug ampoule 48. A hole 49 is provided in the centre of the conical actuator 41. A silicone rubber moulding 50, having a central through-hole 51, substantially fills the space between the actuator 41 and the inner conical surface 44b of the rigid member 43. The rubber moulding 50 extends beyond the conical actuator 41 to form a non-return valve 52 just externally of the actuator 41. A ribbed silicone moulding 53 in the form of a hollow cylinder or collar 53 is provided on the outer face of the actuator 41.

Assembly of the injector 1 is as follows. The tube 46 is bonded to the rigid member 43 and the actuator 41 is bonded to the outer segment 44a of the rigid member 43. The cavity between the rigid member 43 and the actuator 41 is filled by vacuum impregnation with raw silicone rubber such as Dow Corning 734 diluted with silicone oil. A snug fitting wire (not shown) , sharpened to a chisel point, is then placed in the tube 46 such that the sharpened end protrudes sufficiently to form the non-return valve 52.

The silicone rubber 50 is then cured. On demoulding, the sharpened wire is advanced to cut a slit in the rubber moulding, so forming the outlet nozzle and lips 55 of the non-return valve 52. The wire is then withdrawn and the injector 1 is ready for use. The wire may have a polygonal cross-section which will produce for example a star-shape slit, which may have advantages in terms of rapid opening with low rubber momentum. The collar 53 incorporating locating ribs (see below) may be moulded separately and mounted onto the injector 1 with Dow Corning 734 as a separate operation.

When a voltage is applied to the actuator 41, the piezoelectric ceramic increases in thickness and decreases in length. As shown in Fig. 9, the effect is to reduce the volume included between the actuator 41 and the inner conical surface 44b of the rigid member 43. Typically, the thickness of the actuator will be about 0.5 mm which will provide a thickness resonance of 3 MHz. The electrode radius may be 2 mm which will provide a second resonance of 500 kHz. The rubber loading may modify these frequencies slightly, but typically the pressure rise time may be significantly less than a microsecond. On actuation as described, the rubber 50 will be compressed radially inwards, displacing fluid in the fine channel or pump chamber 54 left by the wire used during moulding. If the fluid is air, and as the velocity of sound in air is 0.3 mm/microsecond, air will be pumped out of the valve 52 in preference to transport down the long feed tube 46. The pump may therefore be self-priming. If the fluid is drug solution, the valve 52 is sufficiently close to the pump chamber 54 that the drug solution is rapidly accelerated to skin piercing velocity. Indeed, the convergent flow in the tapered geometry of the non-return valve 52 as shown in Fig. 8 will ensure a high speed low radius front edge to the ejected droplet.

The rubber 50 is compressible. With the bulk modulus of 2 GPa, the 20 MPa drive pressure required to produce a 200 m/s jet will produce a volume rubber compression of 1%. The very thin layer of rubber around the pump chamber 54 will ensure that a minimum of the small pump displacement will be absorbed by rubber compression. Confined rubber becomes very stiff. The effective modulus may increase a thousandfold. The increase in rubber thickness towards the central axis of the injector 1 may be designed to ensure that there is negligible effective rubber stiffening due to enclosure within the pump.

The speed of sound in rubber is approximately 1.4 mm/microsecond. By using a radial rubber displacement geometry with axial fluid flow, the fastest possible pumping arrangement is provided. Typically the maximum distance of the electrodes 42 from the central axis of the injector 1 is of the order of 2 mm so pump strokes of the order of microseconds are possible.

The ratio of the actuator cross-section to that of the pump area may be approximately unity, so quite high pumping pressures are attainable. Typically a maximum of around 20 MPa may be attained. The pump stroke will generate high positive pressure that will propagate in both directions. The outward propagation of pressure will open the lips 55 of the valve 52 and accelerate the ejected drop of liquid. The pressure pulse propagating down the feed tube 46 will be unable to displace the feed fluid significantly because of the inertia of the fluid in the feed tube 46. Provided the pulse is mechanically shorter than the length of the feed tube 46, there will be no effective interruption of feed flow. Because of the excellent acoustic match of the fluid in the feed tube 46 to that in the ampoule 48, there will be no reflected negative pressure wave from the end of the feed tube 46 which might otherwise cause cavitation of the drug within the feed tube 46. The non-return valve 52 is arranged so that there is significant flexibility. The drop in pressure on the pump return stroke will close the valve 52 and suck the lips 55 inwards. The lips 55 will buffer the pump pressure and will provide a continuous low pressure to feed the drug into the pumping chamber 54 which again will prevent cavitation within the drug.

While the conical actuator 41 is a relatively complex shape, it may be formed very straightforwardly by injection moulding PZT powder in an organic binder into a mould tool. Strengthening rims may be formed if desired to preserve tolerances during firing and provide resistance to deformation during operation. The electrodes 42 may be evaporated onto the PZT actuator 41 though a mask. The precision hole 49 at the apex of the conical actuator 41 may be formed with a laser after sintering. This is particularly important as the hole 49 in the actuator 41 effectively defines the output jet diameter.

To obtain uniform standards of operation, the outside electrode 42 may be laser trimmed in real time while measuring the drop volume dispensed under standardised drive conditions. In this manner, very accurate dispensing standards may be maintained relatively simply. The external silicone rubber collar 53 may serve two functions. Firstly, the actuator 41 is a reasonably significant source of ultrasonic power. It is important not to couple this into the patient's skin because of problems of subcutaneous cavitation. The acoustic mismatch between silicone and PZT is such that only 1% of the power is transmitted into the bulk silicone. By fabricating a flexible ribbed collar 53, the coupling from the bulk rubber to the skin may be reduced to an insignificant level. The coupling is further reduced if the collar 53 is mounted just outside the active zone of the actuator. The second function of the collar 53 is to prevent slippage of the injector 1 with respect to the patient's skin. Silicone has a high friction coefficient with most materials, including skin. However, in practice, this can be reduced by lubrication by an alcohol wipe or sweat. If a rubber rib 56 on the collar 53 is applied to the skin 57 as shown in Fig. 11, there will be a pronounced pressure gradient from the tip of the rib 53 to the periphery of contact 58, particularly if the cross-sectional area of the rib 56 increases with increasing distance from the patient's skin as shown. This will effectively pump any lubricant fluid 59 away from the high pressure region, providing a high coefficient of friction along the edge of the rib 53. A number of such ribs 56 radiating from the axis of the system may hold the skin of the recipient in accurate alignment with the jet nozzle. Such a collar 53 may be used with any of the examples described above. While a collar provides additional rigidity from the hoop stresses, a system of free-standing studs may be used instead; such studs may be frustoconical or tapered, for example.

In Fig. 12, there is shown an alternative structure for the pump chamber portion. A hollow conical actuator 41 is again used. However, instead of the rigid member 43 having two conical segments of different angles, the central portion 43b of th. member 4 is spaced from the actuator 41 firstly by a step 60 in the surface of the frame 43 and then by virtue of an increasing taper or increase in conical angle towards its centre. It will be appreciated that other profiles will be acceptable for providing a space between the frame 43 and the actuator 41 for the pump chamber.

Ultrasonic actuators work most efficiently at resonance. However, the amplitude of operation under such conditions depends critically on the quality factor of the resonance, which in turn depends critically on the losses in the system. As these are not well controlled, operation at resonance may lead to erratic dispensation. All examples of the injector 1 of the present invention will operate below the resonant frequency of the actuator 41, so eliminating this problem. The electrical supply may be part of a regenerative circuit so that good electrical efficiency is maintained to permit efficient battery operation.

To obtain large actuator displacements, it is necessary to use soft piezoceramics which have very large loss tangents especially at high field strength and high frequency. However, since the injectors 1 of the present invention will only operate for a maximum of a second in most applications, actuator heating due to the high losses is not a problem. Depoling is not a problem either for, at the high field strengths required, the electrical drive may be in one direction only and repoling will occur with every actuation cycle in practice.

The pulse repetition frequency may be varied to requirement. It seems likely that operation of the injector 1 at similar delivery rates to a hypodermic syringe will minimise the subcutaneous bruising associated with most jet injectors.

The injectors 1 of the present invention may most conveniently be operated with a contact pressure switch.

If the injector 1 is held by an outer casing, the pressure of the actuator tip against the skin may be transmitted to a switch inside the casing and it will be impossible to inject without sufficient contact pressure to inhibit surface sliding. In addition, it will be difficult to shoot the drug accidentally. This may be an important safety feature as the range of the injector in air may be quite significant.

Claims

1. An injector (1) for needle-less administration of a drug, the injector (1) comprising: a rigid member (4,26,43); an actuator (6,22,41); a pump chamber (10,29,54) between the rigid member and the actuator, the pump chamber having an inlet through which a fluid can enter and an outlet through which the fluid can be ejected; and, a non-return outlet valve (12,30,52) at the outlet, whereby operation of the actuator (6,22,41) reduces the volume of the pump chamber thereby ejecting fluid in the pump chamber out through the outlet and non-return outlet valve at a speed sufficient to pierce the skin of a recipient.

2. An injector according to claim 1, wherein the outlet valve (12,30,52) is made of resilient material, the outlet valve being opened by fluid being ejected from the pump chamber on operation of the actuator and closing on removal of fluid pressure or application of negative fluid pressure.

3. An injector according to claim 1 or claim 2, wherein the outlet valve (12,30,52) is formed with at least one slit to provide an outlet nozzle.

4. An injector according any of claims 1 to 3, further comprising a one-way inlet valve (11) at the inlet of the pump chamber (10) for allowing fluid to enter the pump chamber through the inlet and preventing fluid from being ejected through the inlet.

5. An injector according to any of claims l to 4, wherein the actuator (6,22,41) is a piezoelectric, electrostrictive or magnetostrictive actuator.

6. An injector according to any of claims 1 to 5, further comprising a feed tube (2,25,46) connected to the pump chamber inlet.

7. An injector according to any of claims 1 to 6, wherein the actuator (6) forms a wall of the pump chamber (10) , operation of the actuator causing the actuator to extend further into the pump chamber thereby to reduce the volume of the pump chamber.

8. An injector according to any of claims 1 to 7, wherein the actuator (6) is a plate, an edge of the plate forming a wall of the pump chamber (10) .

9. An injector according to any of claims 1 to 6, wherein the pump chamber (29,54) has walls of a resilient material, the actuator (22,41) acting on said resilient walls to reduce the volume of the pump chamber on operation of the actuator.

10. An injector according to claim 9, wherein the actuator is covered by the rigid member (26) , the rigid member having an inner surface opposed to the actuator (22) , the volume between said inner surface and the actuator being substantially filled by resilient material (27) , the pump chamber (29) being formed by a space in said resilient material (27) between the rigid member (26) and the actuator (22) , the actuator extending on operation to force said resilient material against the inner surface of the rigid member (26) and thus into the pump chamber thereby to reduce the volume of the pump chamber.

11. An injector according to claim 10, wherein the inner surface of the rigid member is profiled to minimise the effective rigidity of the resilient material.

12. An injector according to claim 10 or claim 11, wherein the inner surface of the rigid member (26) is frustoconical or hemispherical.

13. An injector according to claim 9, wherein the actuator is a hollow shell (41) .

14. An injector according to claim 13, wherein the actuator (41) is conical, hemispherical or ogival.

15. An injector according to claim 12, wherein the actuator is a hollow shell (41), the rigid member (43) being formed with two concentric segments (44a,44b) , the inner segment (44b) having a profile so as to form a space between the rigid member (43) and the actuator shell (41) , the actuator shell (41) being fixed to the rigid member (43) with the outer segment (44a) of the rigid member being received by the actuator shell thereby forming a space between the inner segment (44b) of the rigid member and the actuator shell, said space being substantially filled by a resilient material (50) , the pump chamber (54) being formed by a space in the resilient material (50) , operation of the actuator (41) causing radially inwards contraction of the actuator shell which acts on said resilient material (50) thereby to reduce the volume of the pump chamber (54) .

16. An injector according to claim 15, wherein the actuator shell (41) is conical and has a taper angle, the outlet (49) being formed at the tip of the cone, the outer segment (44a) being frustoconical and having a taper angle substantially equal to the taper angle of the actuator cone (41).

17. An injector according to claim 16, wherein the inner segment (44b) of the frame (43) is frustoconical having a taper angle greater than that of the outer segment (44a) .

18. An injector according to claim 15, wherein the actuator shell (41) is hemispherical or ogival.

19. An injector according to any of claims 1 to 18, further comprising a projecting collar (53) around the outlet for contacting a patient's skin.

20. An injector according to claim 19, wherein the collar (53) has at least one rib (56) for contacting the recipient's skin.

21. An injector according to any of claims 1 to 19, further comprising at least one tapered projection (56) for contacting the recipient's skin.

22. An injector according to any of claims 1 to 21, further comprising a regenerative circuit for driving the actuator.

23. An injector according to any of claims 1 to 22, wherein the actuator is reciprocating in operation.

24. An injector according to any of claims 1 to 23, comprising means for operating the actuator ultrasonic frequency.