WO1995027449A1

WO1995027449A1 - Cruciate ligament prosthesis

Info

Publication number: WO1995027449A1
Application number: PCT/GB1995/000836
Authority: WO
Inventors: David Martin Pickles; Allen Edward Goodship; Donald James Highgate; Robert Andrew Scott
Original assignee: The Secretary Of State For Health In Her Britannic Majesty's Government Of The United Kingdom Of Great Britain And Northern Ireland
Priority date: 1994-04-12
Filing date: 1995-04-12
Publication date: 1995-10-19
Also published as: EP0749292A1; GB9407237D0; AU2219095A; GB9620916D0; GB2302029A; GB2302029B

Abstract

A bone-fixator is provided for affixing a prosthesis tensile element comprising a fibrous material to a bone, comprising a fixator element capable of being anchored in a bone and an interface for attachment of the tensile element characterised in that the interface has a tensile element engaging element comprising a solid body capable of being gripped by the fixator element such that the fibrous material is trapped between the body and fixator element. Preferably the fixator element and solid body interact such that the fibrous material is uniformly distributed around the solid body. Preferably the fibrous material comprises a sheath into which the solid body locates in use. A ligament prosthesis tensile element is also provided comprising a prestressed hydrogel material and a non-degradable fibrous material; the fibrous material being provided in at least one of the ends of the tensile element and the hydrogel being formed as a rod which expands to prevent the fibrous material from collapsing toward the element longitudinal axis under physiological load.

Description

CRΠCTATE I.TΠAMENT PROSTHESIS.

The present invention relates to novel ligament prostheses, and bone-fixators allowing improved function of such prostheses; particularly relating to anterior cruciate ligament (ACL) prostheses.

Although there are number of anterior cruciate ligament prostheses in current use, none has performed sufficiently well to gain general acceptance. No prostheses have yet received FDA approval for routine use in humans, largely due to the unacceptably high incidence of mechanical failure and adverse reactions to particulate debris arising from fragmentation. Failures may largely be attributed to poor prosthesis designs which fail to take into account the biomechanical properties of the natural ligament. Typically failure is of the fixation device or due to mid-substance rupture.

It is required that the biomechanical properties of the prosthesis should satisfactorily match those of the natural ACL, in the following respects: (i) sufficient overall tensile strength; (ii) satisfactory mimicking of force/deformation profile low and high modulus regions; (iii) minimal hysteresis; (iv) minimal creep; (v) excellent elastic recoverability.

The inventors have paid special attention to measuring ACL strain in the walking animal using implanted displacement transducers to determine the physiological strain environment of the ligament. In common with other soft tissues, the force deformation profile of the ACL shows a characteristic toe region where the modulus starts low and increases to a linear, relatively high value. The toe region of the ACL may have an important role in defining the complex motion of the knee joint, which shows gliding of the femoral condyles over the tibial condyles as the joint rotates and the leg extends.

The inventors have found that use of prostheses based on synthetic braids alone does not provide particularly good matching of the 'toe¹ region. In order to achieve the two domains of markedly differing mechanical properties they have then turned their attention to choice of a composite prosthesis design, using correspondingly different materials.

Furthermore they have found that when impregnation of braid with silicon rubber is undertaken a compliant, compressible prosthesis element is provided. When several different configurations were subjected to tensile testing, the force-deformation curves of all species were indistinguishable from ordinary, non-impregnated braids, however this was shown to have no effect on the gross tensile properties of the braid.

The present inventors have now accurately quantified the properties of the ACL and have succeeded in developing a biomechanically-matched prosthesis which aims to prevent prosthesis wear and to maintain a physiological gauge length; these being features which no other conventional system can achieve. The prostheses may be used for replacement of the natural ligament in animals in general, with the dimensions and strengths of materials matched to its weight or particular requirements, but is particularly intended for human use.

The inventors have found that in order to achieve the improvement required it is possible to use a continuous, prestressed hydrophilic element in conjunction with a synthetic braid whereby a radically new approach to the problem of fixation and initial tensioning of the ACL during surgical implantation is provided.

Hydrophilic polymeric materials are characterized by the ability to take in large amounts of water and reach a state of stable hydration without suffering solution or long term degradation. Equilibrium water contents of 10 - 98% by wet weight may be readily achieved. During absorption of water hydrophilic material normally swells isotropically and changes from a hard, rigid structure, becoming soft and elastic. It may be accurately machined and polished while dry, and is subsequently hydrated to become soft, conforming and biocompatible. The linear expansion ratio (Xwet/Xdry) , which lies typically between 1.1 and 3>5t may be accurately controlled. Hydrophilic polymers have been used extensively in contact lens applications, and also in stoma care dressings which necessitate subcutaneous implantation for biocompatibility assessment. More recently, hydrogels have been used as dusting powder in surgical gloves. There have been no reports of adverse reactions.

Traditionally, hydrophilic materials (such as those used for contact lenses) have a low 'notch-tear¹ strength. However, interpenetrating systems, particularly those based upon acrylonitrile polymers, which combine good elasticity with high notch tear strength can be ¹prestressed' when dry so that on hydration the system may contract in one or two dimensions. Thus the axial length of a dry hydrogel rod may be arranged to either increase, remain the same, or decrease by any contraction ratio up to 5∑1ϊ in the latter case there is of course a lateral expansion due to the hydration of the material and the axial expansion.

It is self-evident that no hydrophilic material could be used as a stand-alone prosthesis, since all such substances are too elastic to provide the necessary tensile properties. However, it may be used in two ways to generate a low modulus region to a prosthesis: firstly, as a core inserted inside an external sheath which would provide the required strength, or secondly, as a continuous metric containing a dispersion of reinforcing fibres. In the latter instance the hydrogel needs to be cast with the appropriate fibre elements in situ. In both cases the hydrogel would act in compression resisting the tendency for the prosthesis tensile component to collapse inwards under axial load.

The principle of using a composite low-modulus gel core and a load -bearing outer sheath to match the biomechanical properties of a natural ligament or tendon has already been established (Dore & Drouin, US 4301551, and Shah, 1987). Hydrogels have also been used in an experimental prosthesis using PHEMA hydrogel reinforced with degradable poly(lactic acid) fibres (Kolarik et al.(198l) J. Biomed. Mat. Res 15; Davis et al., 1991)- Unlike these prostheses, the present invention provides a low modulus core which expands upon rehydration to fulfil its role.

The use of hydrophilic polymer material allows a prosthesis to be manufactured chemically complete before implantation but that changes shape permanently in a controllable and reproducible fashion upon exposure to tissue fluids. Furthermore experience with soft contact lenses has shown the materials to be extremely mechanically stable in fatigue testing.

Thus a first aspect of the present invention provides a ligament prosthesis tensile element comprising a prestressed hydrogel material and a non-degradable fibrous material; the fibrous material being provided in at least one of the ends of the tensile element and the hydrogel being that which expands to prevent the fibrous material from collapsing toward the element longitudinal axis under physiological load.

Preferably the fibrous material runs the entire length of the tensile element and is provided as a sheath, or as fibres embedded in the hydrogel. Preferably the hydrogel expands more axially than radially.

Preferably the prosthesis tensile element has a tubular outer sheath, preferably a biocompatible braid, in which the tensile element engaging element is positioned in use. The braid is conveniently any biocompatible material capable of load bearing, such as non -impregnated or impregnated polymeric braid. One convenient material is polyester braid in the form of a cord. Such braid cords may be provided in a variety of strengths and sizes and are readily available form such sources as English Braids Ltd, Spring Lane, Malvern, Worcestershire, WR14 1AL United Kingdom.

The present inventors have determined that by varying primarily the diameter of the hydrogel core, and secondarily, its nature (i.e. chemical species/water content etc) the mechanical properties of the hydrogel-braid composite can be made to match those of the natural structure. The interrelationship of braid lay angle and core diameter has been demonstrated; use of a prestressed hydrogel core means that the lay angle is recoverable following a load cycle.

The braid should biomechanically match the ligament it is to replace. For example for a sheep ACL an English Braids Ltd R27K braid cord may be used having 2.7mm diameter with yarns to a total of 1500 denier or metric equivalent of 1550 decitex. These cords are sold with a centre core which is added to maintain roundness, and this is removed and filled with hydrogel when preparing the prosthesis tensile element.

In the preferred embodiment the braid is filled with a hydrogel core that expands on further hydration with water such as to give the tensile element a full resilient nature. Many suitable hydrogels will be known to those skilled in the art. Preferred such gels have linear expansion ratios of 1.0:1.0 to 1.8:1, with gels of ultimate water content between 20 and 80%. More preferably the linear expansion ratios of is around 1.2:1.0 and the ultimate water content is around 30 . Hydrogel cores are conveniently formed as rods and machined in unhydrated form down to desired diameter for threading down the braid. For the sheep prosthesis this is about 13 mm. Rehydration is achieved eg. by immersion in isotonic saline prior to testing but can be rehydrated to below its ultimate water content prior to implantation.

In a most preferred embodiment of the present invention the prosthesis tensile element is provided with a further sheath, also of biocompatible material, positioned around the prosthesis tensile element such that the tensile element can slide therein. Preferably this further sheath is such that it allows, or even stimulates, growth of tissue around it such as to fix the prosthesis in place with time.

In order to use tensile elements containing fibrous materials and polymeric hydrogel materials in a prosthesis assembly it is necessary to provide a bone fixator/prosthesis interface capable of holding such a prosthesis in place. Thus in order to allow use of the prostheses of the present invention the inventors have provided a suitable novel interface.

In order to provide a satisfactory prosthesis design the interface should be designed to allow prosthesis tensile elements, particularly those provided with such a prestressed core, to act in a manner closely matched to natural ligaments and should further permit satisfactory prosthesis tensioning to be applied at implantation and preferably allowing retensioning in the future should the need arise. Furthermore, the overall prosthesis length will be largely dependent on the geometry of a given knee joint and will not be known prior to the operation. Such variation must be accommodated into the design. The gap between the end of the hydrogel and the fixator should be filled appropriately to avoid discontinuity in the material properties of the whole structure. The length of the gap will relate to the overall length.

Currently marketed prostheses use either screws, bollards, toggles, staples or combinations thereof and are fundamentally flawed in that they result in a prosthesis of an undefined gauge length, considerably longer than that of the natural ACL.

The present invention seeks to anchor the substance of the prosthesis to the bone using a a fixator element capable of being anchored in a bone, and an interface for attachment of the tensile element characterised in that the interface has a tensile element engaging element comprising a solid body capable of being gripped by the fixator element such that the fibrous material of the prosthesis is trapped between the body and the fixator element.

This means that load-bearing in the prosthesis tensile element commences close to the articular surfaces closely approximating the length of the natural ligament. The system thereby simulates the physiological origin and insertion of the ACL, and also provides for a gradual transfer of load from the tensile element to the fixator, avoiding stress raisers which would render the element prone to fatigue failure. Since there is of necessity a region between the engaging element and the hydrogel core where the braid is not fully internally supported, this region is vulnerable to unrecoverability (braid lay angle reconfiguration) . The problem can be addressed in this design by interposing an axially prestressed hydrogel section between the engaging element and the hydrogel core end, designed to expand axially upon hydration but maintaining its diameter equal to that of the main hydrogel core, thereby preserving the braid configuration. These sections act to fill any discontinuities which may manifest themselves following repeated load cycles.

The sections may take the form either of small shims of different lengths to be selected by the surgeon at implantation and are intended to be inserted down the ends of the braid during assembly, or as a portion of the material embedded into a hollow cavity at the tip of the engaging element. The sections should preferably be of the diameter of the tensile element within the sheath but expand axially on hydration, in vivo or in vitro, to take up slack.

For the purpose of allowing the length of the tensile element to be altered the hydrogel shims may be provided such that they can be inserted into the tensile element sheath by the surgeon installing the prosthesis as required in order to take up any slack in the braid. It may be desired that such shims, or the core itself, only expand in particular directions, eg. axially rather than radially. Such capability is provided by use of hydrogels such as those described in GB 1566552 (Highgate) . In the further option where the engaging element incorporates a hydrogel insert this is similarly biased such that it expands only in one direction on further hydration when in situ (ie. in the animal or patient). This gel may be formed with a number of hollows therein in order that expansion within the element is avoided or reduced.

In a second aspect of the present invention there is provided a bone-fixator for affixing a prosthesis tensile element comprising a fibrous material to a bone, comprising a fixator element capable of being anchored in a bone, and an interface for attachment of the tensile element characterised in that the interface has a tensile element engaging element comprising a solid body capable of being gripped by the fixator element such that the fibrous material is trapped between the body and the fixator element.

Preferably in use the fibrous material is uniformly distributed around one or more cross sections of the solid body, and most preferably is provided as a sheath into which the solid body locates in use. In this most preferred form the sheath containing the body is inserted into a bore in the fixator element and the portion of the sheath around the body is firmly gripped between them to lock in position. The fixator element may already be fixed into a bone or fixed later.

Preferably the solid body is formed to have a width that increases from that which will fit into the unexpended braid to that which is greater in width than the unexpended braid diameter and preferably then reduces again, most preferably with conical or parabolic shaped axial ends. This is preferably a bullet-like element with concentric interference-fitting conical faces (an axial bullet) such that it can be driven into firm interference fit with the fixator element by use of a force sensitive tool. Alternatively the body may for example be locked in place using a locking element on the fixator, such as a locking nut.

Use of a such a solid body allows the tensile element length to be adjusted axially by inserting it through a bore in the fixator element, inserting the engaging element, eg. bullet shaped body, into the end of the tensile element such that it is sheathed thereby, adjusting the length of tensile element to that required and then securing the engaging element into engagment with the fixator element.

In current devices tension in the prosthesis must be selected by the surgeon at implantation, after which it cannot be adjusted. In contrast when using a device according to the present invention the engaging element, eg. bullet body, may be retracted, and braid drawn back through the fixator to the appropriate tension, and the element engaged again.

In a preferred form of the invention, the fixator comprises an inner fixator element which can moved within an outer bone-engaging element (eg. a self tapping hollow bone screw) . Preferably the inner element is slidably mounted within the outer bone engaging element and can be retracted into it or extended out from it by means of a threaded collar thereby allowing fine-tuning of the tension of the tensile element once the device is in its final position. Preferably relative rotation between the fixator element and the bone-engaging element is limited by means of a cooperating lug/groove or other rotation-resisting arrangement.

In addition to the advantage of allowing precise tensioning, the use of a separate inner fixator element which can be located within an outer bone-engaging element permits the preassembly of a readily fixatable tensile unit before surgery. Thus an appropriate length of tensile element may be secured at each end to an inner element fixator. A set of several such preassembled units of different lengths is supplied to a surgeon who can, during the operation, select the most appropriate for his patient. The unit can then be threaded through bone screws after they have been implanted, and fine tuning of the tension can then be carried out using the threaded collar.

Thus a third aspect of the invention comprises such a unit comprising a suitable fibrous tensile element attached to one or two inner fixator elements as described above.

In the event of a catastrophic failure of the braid in vivo, the devices of the present invention would be far more suitable for a revision operation than any existing prosthesis, since the bone fixators would be left in situ and a replacement braid/hydrogel unit could be fitted without further interference to the bone. Alternatively, in embodiments incorporating a separate inner fixator element within an outer bone engaging element, the tensile unit (i.e. tensile element plus one or both inner fixators) could be changed.

Further provided by the present invention is a prosthesis and fixator assembly comprising a tensile element, a fixator element and a tensile element engaging element; all as described above.

The prosthesis and prosthesis/fixator interface of the invention will now be described further by reference to the following non-limiting Examples and Figures by way of illustration only. Further embodiments of the invention will occur to those skilled in the art in the light of these.

FIGURES.

Figure 1: shows a typical force deformation curve for collagenous tissue of a sheep anterior cruciate ligament.

Figure 2: shows force strain curves for number of hydrogel cored braids as described in Table 1 as compared to a natural ACL.

Figure 3: shows a plot of peak force achieved during application of 10% uniaxial strain to specimens described in Table 1.

Figure 4: shows a plot of normalised stiffness of the linear elastic portion of the force strain curve during application of 10% uniaxial strain v log (cycle number) for each specimen of Table 1.

Figure 5- shows a plot of normalised energy loss (hysteresis) against log (cycle number) during tensile testing for the specimens of Table 1.

Figure 6: shows a plot of braid lay angle as a function of applied load for the specimens of Table 2.

Figure 7- shows a schematic diagram of a bone fixator/prosthesis tensile element interface of the present invention in end view (a) and a longitudinal cross section (b) , the prosthesis engaging element being held in a locked condition by interference fit between its surfaces and the fixator provided by impacting the two together under a measured torque.

Figure 8: shows a schematic cross section of a unit according to the present invention comprising a prosthesis tensile element wherein each end is attached to a fixator body in readiness for connection to a bone screw.

Figure 9: shows a cross section through an interface as described in Figure 7 but wherein a hydrogel core element is incorporated in the end of the engaging element for the purpose of taking up slack.

EXAMPLES.

EXAMPLE 1;—Properties of prosthesis tensile element cores.

In all cases the prosthesis were modelled upon the sheep ACL and were made using English Braids Ltd polyester braid R27K; this likely to need replacement with similar but stronger braid for use in humans. The specification of R27K is as follows:

R27K: 2.7mm diameter multifilament polyester braided cord provided with a central core. The cover is braided using a 16 carrier machine, 8 going cw and 8 going acw. Each carrier has a yarn or yarns to the total of 1500 denier or the metric equivalent of 1550 decitex; usually 1 x 1500 or alternatively 1 x 1000 plus 1 x 500 denier. The polyester yarn from which this and other English Braids Ltd products is of 1000 denier (1100 decitex), 1117 Grams per 10000m with a breaking load of 89.2N, tenacity 79«9cN/tex, extension at break 13.8%, tensile shrinkage at 180°C of 14.1% and oil content 0.6%. A number of hydrogel core/braid sheathed tensile elements were prepared having properties as set out in Table 1. Hydrogel used was

Table 1 Description of tested specimens

Specimen Core Hydrogel Hydrogel Gauge

Ref Type OD (dry) OD (mm) length (mm)

(mm) (hydrated) (hydrated) (calc. ) (measured) (mm)

PR37 none n/a n/a 200 ('Empty' braid)

PR53 ED4 1.20 1.98 116

PR44 ED4 1.32 2.18 102

PR39 ED4 1.82 3.00 125

ACL none n/a n/a 24 (Sheep ACL)

designated ED4 and had an ultimate water content of 75% with a linear expansion ratio of 1.65:1. The gel controls the angle of lay of the braid and thus provides the non-linear load/extension characteristics desired in the finished unit. This interaction is a substantial simplification of the complex process described by Dore in US 4301551 which also uses core/braid interaction.

In Figure 1 there is shown a typical force deformation curve for collagenous tissue wherein a uniaxial tensile test of a sheep anterior cruciate ligament is shown. The specimen was well conditioned by repeated cycling, and the origin of the curve (zero strain, zero load) is defined as the point where load is just detectable, determined by repetitive fitting of cubic polynomial, Riemersa and Van den Bogart, Acta Anatomica, (1992). A regression line has been fitted to the linear portion of the curve, defining the toe and linear elastic regions of the force-strain curve.

In Figure 2 force-strain curves for specimens of different hydrogel cores inside R27K braid (detailed in Table 1) are compared to a natural sheep ACL. Rods of hydrophilic material were machined in their unhydrated form down to the indicated diameters, threaded down the centre of a 300mm length of braid, and immersed overnight in isotonic saline to achieve hydration of the polymer. Each specimen was then attached to a Dartec materials testing machine, gripping the braid at each end with a simple parallel clamp immediately adjacent to the end of its hydrogel core. An X-Y table attached to the actuator ensured accuracy of the axial alignment of the sample. The specimen was then stretched until tho load cell recorded a force of 1-2N, to define the gauge length. Tests then consisted of cycling to a 10% strain at a rate of 2%/sec. Following a single conditioning cycle, after which the machine was re-zeroed, the specimen was cycled 111 times: force-displacement data was recorded on a PC at 50Hz for cycles 1-11 and 31, 4l, 51. 61, 71. 81, 91. 101 and 111. Comparable experiments were also carried out for 'empty' R27K braid, and for a sheep anterior cruciate ligament in vitro, still attached to femur and tibia and mounted in special clamps to permit uniaxial tensile testing. The complete force-strain curve (ie applying and removing tension) is shown.

Figure 3 shows the peak force achieved during application of 10% uniaxial strain to each specimen described in Table 1, plotted against log(cycle number) . The final point on each line corresponds to the cycle plotted in Fig 2.

Figure 4 shows the stiffness of the linear elastic portion of the force-strain curve during application of 10% uniaxial strain to each specimen described in Table 1 plotted against against log(cycle). number) . Stiffness was evaluated as the slope of a line drawn between the 100N and 240N data points, and normalising this for variation in specimen length by multiplying by gauge length. Stiffness was not measured for specimen PR39 since the linear elastic region of the curve was not attained within the test.

Figure 5 shows the energy loss (ie. hysteresis) during tensile testing of each specimen described in Table 1, plotted against log(cycle number) . Hysteresis was evaluated by measuring the difference in the areas under the tensile and relaxing phases of each cycle, using a computer algorithm.

EXAMPLE 2: Bra-it- lav anglp nf prosthesis tensile element cores.

A number of hydrogel core/braid sheathed tensile elements were prepared having properties as set out in Table 2. Hydrogel used was

Table 2 Description of tested specimens

Specimen Core Hydrogel Hydrogel Ref Type OD (dry) OD (mm) (mm) (hydrated) (calc. )

A none n/a n/a B D/CG/11 1.81 2.0 C D/CG/11 1.96 2.2 D D/CG/11 2.68 3.0

designated D/CG/11 and had an ultimate water content of 34% with a linear expansion ratio of 1.12:1.

Figure 6 demonstrates how the configuration of a braid is precisely controlled by the diameter of the hydrogel core within. Each specimen braid of Table 2 was loaded and unloaded experimentally and the angle subtended by its constituent fibres to the loading axis was monitored using the crosshairs of a travelling microscope. As can be seen, the initial braid lay angle is dependent on the nature of the hydrogel insert and the recoverability of initial braid lay angle (an important feature in a braid prosthesis) is much improved in the hydrogel containing the large diameter hydrogel core.

EXAMPLE 3; Bone fixator/prosthesis tensile element interface Figure 7 shows a schematic diagram of a bone fixator/prosthesis tensile element interface in end view (a) and a longitudinal cross section (b). The prosthesis engaging element consists of a steel cone body (1) within the bore of a braided sleeve (2), fitted inside a steel bone fixator (3).

The fixator (3) is slidable mounted within a bone screw (4) and is secured therein by a collar (5). Rotation between the fixator and the bone screw is limited by slots in the fixator (6) one of which cooperates with a lug (7) on the inner surface of the bone screw. For precision tensioning the fixator can be longitudinally relocated within the bone screw by means of a mutual thread between the fixator and the collar.

Tensioning may also be achieved by moving the braid within the fixator. In either case the braid is locked in position by torquing down the steel cone body using tool (8) to (eg. using around 1500 N controlled force) provide an interference fit between it and the fixator with the braid between the two.

Figure 8 shows a schematic cross section of a tensile unit comprising a braid (2) wherein each end is secured to a fixator (3) by a torqued down steel cone body (1).

In a further embodiment, Figure 9 shows an engaging element and tool (8) in place where the element body tip is hollow as shown, open at its narrow end, and filled with dehydrated, prestressed hydrophilic insert (9). On hydration, the insert expands out of the cone to fill the gap between the core in the braid (2) and body (1). In the modification shown inclusion of known and controlled void spaces (10) in the hydrophilic material allows the pressure exerted on the end of the core after hydration may be adequately controlled. This concept is known with regard to a patented design of tooth root filling.

EXAMPLE 4; Implantation of the bone screw

In use the bone screw (4) is implanted as follows. The surgeon identifies the femoral or tibial fixation site and drills a pilot hole using a guide wire. Leaving that in place, a larger hole is bored towards the joint centre using a stepped reamer: this leaves external holes in the tibia and femoral condyle large eneough to accept the bone screw and internal holes in the articular surface just large enough to accept the fixator. The bone screw is inserted using a custom wrench which engages the hexagonal top of the screw and terminates in a round rod of similar diameter to the fixator which temporarily engages in the smaller hole in the articular surface to maintain concentricity as the bone screw is driven home. When tight, the base of the bone screw abutts against a shelf of cortical bone approximately 4 mm thick. A fixator, or fixator/tensile element unit, is then secured to the bone screw such that the end of the fixator is flush with the articular surface, thereby preventing the braid from abrainding against the sharp exit of the bone tunnel and maintaining the active length of the prosthesis as close as possible to that of the natural ligament.

Claims

CLAIMS .

1. A ligament prosthesis tensile element comprising a prestressed hydrogel material and a non-degradable fibrous material; the fibrous material being provided in at least one of the ends of the tensile element and the hydrogel being formed as a rod which expands to prevent the fibrous material from collapsing toward the element longitudinal axis under physiological load.

2. A tensile element as claimed in claim 1 wherein the fibrous material runs the entire length of the tensile element.

3. A tensile element as claimed in claim 1 or 2 wherein the fibrous material is provided as a sheath around the hydrogel.

4. A tensile element as claimed in any one of claims 1 to 3 comprising fibres embedded in the hydrogel.

5. A tensile element as claimed in any one of claims 1 to 4 wherein the hydrogel expands more axially than radially on rehydration.

6. A tensile element as claimed in any one of claims 1 to 5 wherein the fibrous material comprises a biocompatible tubular outer sheath of braided fibres.

7. A tensile element as claimed in claim 6 wherein the braided fibres comprise a polymeric material.

8. A tensile element as claimed in claim 7 wherein the polymeric material is a polyester.

9. A tensile element as claimed in any one of claims 1 to 8 provided with one or more axially prestressed hydrogel sections adapted to be inserted between a fixator/tensile element interface and the hydrogel of the tensile element, the sections being capable of expanding axially upon hydration to take up slack ano attaining a diameter substantially equal to that of the hydrogel of the tensile element.

10. A tensile element as claimed in claim 9 wherein the hydrogel sections take the form of shims of different lengths.

11. A bone-fixator for affixing a prosthesis tensile element comprising a fibrous material to a bone, comprising a fixator element capable of being anchored in a bone and an interface for attachment of the tensile element characterised in that the interface has a tensile element engaging element comprising a solid body capable of being gripped by the fixator element such that the fibrous material is trapped between the body and fixator element.

12. A bone-fixator as claimed in claim 11 wherein the fixator element and solid body interact such that the fibrous material is uniformly distributed around one or more cross sections of the solid body.

13« A bone-fixator as claimed in claim 11 or 12 wherein the fibrous material comprises a sheath into which the solid body locates in use.

14. A bone-fixator as claimed in claim 13 wherein the fixator element has a bore adapted to receive the sheath containing the body whereby the portion of the sheath around the body is gripped between the body and fixator bore such as to locked in position.

15- A bone-fixator as claimed in any one of claims 11 to 13 wherein the solid body has a width that increases from that which will fit into the unexpended braid to that which is greater in width than the unexpended breid diameter.

16. A bone-fixetor es cleimed in claim 15 wherein the width of the solid body reduces again after increasing.

17. A bone-fixator as claimed in claim 15 or 16 comprising one or more conicel ends in the axial direction of the bore.

18. A bone-fixator as cleimed in claim 15 wherein the solid-body has an end having concentric interference-fitting conical or parabolic profiled ends with diameters such thet it can be driven into firm interference fit with the fixator element.

19. A bone-fixator es claimed in claim 18 wherein the solid body is adapted to be installed into the fixator by use of a force sensitive tool.

20. A bone-fixator es cleimed in eny one of clβims 15 to 19 wherein the solid body is locked in piece using e locking element on the fixetor.

21. A bone-fixetor es cleimed in any one of claims 11 to 20 wherein the tensile element length may be adjusted axially by inserting it through a bore in the fixetor element with the solid body inserted into the end of the sheeth such es to be sheathed thereby, adjusting the length of tensile element to that required and then securing the solid body into locked engagement with the fixator element through the sheath.

22. A bone-fixator as claimed in eny one of cleims 11 to 21 wherein the fixetor element includes en inner fixetor element which can be moved within an outer bone-engaging element.

23. A bone-fixator es cleimed in cleim 22 wherein the inner fixator element is slidably mounted within the outer-bone engaging element.

24. A bone-fixator as cleimed in cleim 22 or 23 wherein the inner fixator element can be moved within the outer bone-engaging element by means of a threaded coller such es to be eble to fine-tune the tension of the tensile element once it is instelled in its final use position.

25. A bone-fixator as claimed in any one of claims 11 to 24 wherein the solid-body has a portion of hydrogel embedded into a hollow cavity in its tip facing the tensile element in use whereby the portion of hydrogel expands axielly on hydretion.

26. A fixeteble tensile unit comprising e fibrous tensile element etteched to one or two inner fixetor elements es described in cleims 23 to 25.

27. A bone-fixβtor/prosthesis tensile element essembly comprising e bone-fixetor element as claimed in any one of claims 11 to 26 and a prosthesis tensile element as claimed in any one of claims 1 to 10.

28. A prosthesis tensile element, bone-fixator, or assembly comprising these, as described in the Examples 1 to 3 herein, as cleimed in cleim 1, 11 or 27.