Since the early 1970s, total ankle arthroplasty has been considered for the treatment of severe erosions of the articular surfaces of the ankle in humans.14 In the following decade, the disappointing results of long-term followup clinical studies22,25,44 of the pioneering designs has left ankle arthrodesis as the surgical treatment of choice for these patients.20 The evident high incidence of nonunion, secondary degenerative changes at neighboring joints, a high incidence of postoperative infection, and total loss of motion resulting from this surgery3,9 have contributed to a remarkable renewed interest in total ankle arthroplasty during the past decade.7,10,15,17,18,39,45,46
Although more recent clinical reports are a little more satisfactory,1,4,23,38,42,51 total ankle arthroplasty still is not as successful as total hip and total knee arthroplasties. For the total ankle arthroplasty to be considered a valuable alternative to ankle arthrodesis, an effective range of ankle mobility must be recovered. The disappointing clinical results of the current generation of total ankle arthroplasty has been related to poor understanding of the structures guiding joint mobility.2,45 In devising joint replacements in humans, most attention has been given to the geometry of the prosthesis components in relation to the morphologic features of the intact articular surfaces.12,35,36,40,43,47 Little attention has been given to the restoration of the complementary roles of the ligaments and articular surfaces.16,19,34 Particularly for the ankle, replication of natural anatomic shapes seems to be the only guideline for prosthesis design,6,21 with the even more questionable solution of replacing only one of the two articulating surfaces with natural shapes.5,24
Numerous investigations8,15,26–33 have shown how all the passive structures of the ankle control and limit joint motion. Experiments30,32,33 have described movements at the ankle and subtalar levels in virtually unloaded conditions. Two-dimensional mathematical models of the intact ankle28,29 have shown that a preferred path of complex motion is guided by the articular surfaces and the ligaments interacting together in a complementary manner. These findings about the mobility of the intact ankle in the sagittal plane have been supported by a preliminary study on joint stability with a three-dimensional arrangement of the ligaments.8 Although limited to the sagittal plane, these models have contributed to the design of a possible new total ankle implant.26,27
Based on this body of research, we have developed a new total ankle prosthesis that is compatible with the physiologic mobility and with the natural role played by the ligaments. The purposes of this study were to provide an understanding of how our previous investigations led to the design, to describe salient features of the design, and to report on our preliminary observations of ankles replaced with the new total ankle prosthesis.
Conventional and Novel Criteria for the Design of Total Ankle Replacements
Conventional Considerations for Total Ankle Arthroplasty Design
In general, human joint replacement design is subject to anatomically based and functionally based expectations, mechanical and material resistance analysis, surgical and clinical requirements, and industrial and manufacturing constraints. For total ankle arthroplasty, specific issues remain: (1) restoration of physiologic motion and normal ROM; (2) injury resistance particularly in inversion and eversion trauma; (3) preventing talar component subsidence with minimum talus resection; and (4) feasibility of implantation, because of the small dimensions of the joint. Because the subtalar joint often is damaged, subtalar motion also should be restored in the replaced ankle, particularly allowing physiologic rotations of the ankle complex in the transverse (internal and external rotation) and frontal (inversion and eversion) planes.
As with other joint replacements, the traditional dilemma between mobility and congruency must be addressed.16 Unconstrained or semiconstrained designs allow the necessary mobility but require incongruent contact, thereby giving rise to large contact stresses and potentially high wear rates. Conversely, congruent designs produce large contact areas with low contact stresses but transmit undesirable constraint forces that can overload the fixation system at the bone-component interface.
These theoretical assumptions have been proven in practice for the knee, where the least conforming designs have had the worst features of destructive wear.11,49 Mobile meniscal-bearing elements have been adopted in some designs to allow for complete congruence and minimally constrained components over the entire ROM, but with the associated risk of subluxation or dislocation or both.41
Recent Findings on the Mobility of the Ankle Complex
Mobility in joints in humans may be understood by study of the role played by the passive joint structures in virtually unloaded conditions.30,50 It has proved helpful to study the geometry of the articular surfaces and of ligaments in all possible joint configurations during passive motion (mobility) before moving on to study of the response to external loads (stability). Equivalently, during implantation of joint prostheses, the surgeon usually checks for the restoration of acceptable joint mobility before any assessment of joint stability.
Experiments done on intact specimens from patients who had had below-knee amputation30,32,33 have been interpreted by mathematical models.8,28,29 In virtually unloaded conditions, motion of the ankle complex (calcaneus with respect to the tibia and fibula segment) has one-degree-of-freedom. It occurs mostly in the sagittal plane, and at the ankle (tibiotalar) level, with very little motion at the subtalar (talocalcaneal) level. Our model suggests that this motion is guided by the nearly isometric rotation of one fiber within the calcaneofibular ligament on the lateral aspect of the ankle, and one fiber within the tibiocalcaneal part of the deltoid ligament. These two fibers rotate isometrically about their origins and insertions, whereas other fibers of these ligaments slacken as soon as the ankle starts to plantar flex. Ligaments located more anteriorly slacken during dorsiflexion; those located posteriorly slacken during plantar flexion. The instantaneous center of rotation lies at the intersection of the isometric fibers. The instantaneous center of rotation and the point of articular contact moves from back to front relative to both bony segments during dorsiflexion; therefore, rolling and sliding occurs. Because of the changing positions of the instantaneous center of rotation and the muscle lines of action, the lever arm lengths change with flexion. Flexor muscle lever arms are maximized in dorsiflexion, extensor muscle lever arms are maximized in plantar flexion, optimizing the mechanical advantage of the muscle forces during level walking.28
Compatibility between the articular surfaces and ligaments over the flexion arc is ensured if the common normal (the line mutually perpendicular) to those surfaces at their point of contact passes through the instantaneous center of rotation.16 This allows a considerable flexibility in total ankle arthroplasty design because, for every choice of the shape and location of, say, the tibial component, the compatible shape of the talar component can be calculated from the common normal theorem.
Features in the Sagittal Plane of a Possible New Ankle Replacement
More recent studies26,27 were aimed at designing the general type of articulation in the sagittal plane of a possible new ankle replacement, which could guarantee compatible function between prosthetic surfaces and isometric ligament fibers, while possibly maintaining large congruity in all joint positions. Several ligament-compatible pairs of articular surfaces were tested using the original four-bar linkage model.29 The main features of the ankle replaced by nonconforming two-component and by fully conforming three-component designs with either flat, concave, or convex tibial surfaces were assessed.
A three-component prosthesis was favored because complete congruence over the entire motion arc is attainable. Among the infinite range of plausible pairs of tibial and talar articular surfaces, all compatible with physiologic ligament tensioning, a convex-tibia fully congruent three-component prosthesis was favored eventually because of the better degree of entrapment of the meniscal bearing with associated smaller risk of dislocation (Fig 1). The tibial arc radius was selected as to allow for an entrapment similar to that of previous designs,5,24 while ensuring a clinically acceptable thickness of the tibial component. The AP length of the meniscal bearing was optimized to maximize the associated level of entrapment and the contact area, limited only by the risk of impingement or overlapping of the components at the extremes of the flexion arc.
The prosthetic articular surfaces are constrained minimally to enable the soft tissues to control the physiologic motion at the joint. This feature allows replication of almost exactly the original pattern of ligament fiber slackening and tightening and motion of the instantaneous center of rotation.26,27 This simulation work also revealed that the talar sagittal surface compatible with ligament isometry and a flat tibial surface has a larger radius of curvature than that of the natural talus, unlike some current designs.
Based on these prior investigations, we have developed a new total ankle implant. The interface between the tibial and meniscal components is a spherical joint (Fig 2), allowing rotation about each of three axes. The articular surface of the tibial component is a convex segment of a sphere. The selected radius of curvature for the arc in the sagittal plane (RTi) is replicated for the arc in the frontal plane. The upper surface of the talar component shows a circular convex arc in the sagittal plane (radius RTa), designed to be compatible with ligament isometricity, with a concave sulcus when seen in the frontal plane. The meniscal-bearing component has concave upper and lower surfaces, fully conforming with the corresponding tibial and talar surfaces.
More design features, including fixation elements, are shown in Figure 3. The tibial component has a lower surface shaped as a convex spherical segment whose center should rest just anterior to the midline of the prepared tibial surface. Similar to previous meniscal-bearing designs, tibial component fixation is achieved through two parallel cylindrical bars running anteroposteriorly on the upper flat surface. The upper talar articulating surface is saddle-shaped. It is a surface of revolution, generated by rotating a concave curve about a mediolateral fixed axis. The radius of this arc in the sagittal plane is, however, different from that of the tibial component, and calculated to be compatible with the isometric rotation of the guiding ankle ligaments. When the prosthesis rests in neutral position, the sagittal arc is slightly longer posteriorly because the motion expected in plantar flexion is larger than that expected in dorsiflexion. Similar to current total ankle arthroplasty, the talar component is narrower posteriorly than anteriorly to recreate the natural talar anatomy and diminish the risk of posterior component impingement. The talar component rests on a central horizontal flat surface and on anterior and posterior chamfers made on the prepared talar dome. On the undersurface, two pegs are used for fixation, one on the anterior chamfer, located more to the right, and the second in the central surface, located more to the left. The pegs are both oriented posteriorly to facilitate component implantation.
The articulating surfaces of the tibial and talar components are highly polished. To facilitate fixation of both these components with the bone, porous coating with hydroxyapatite is used on the undersurfaces.
A PE meniscal-bearing component is designed to have concave upper and lower surfaces fully congruent with the corresponding tibial and talar prosthesis components irrespective of joint position. It is slightly longer posteriorly, particularly the upper surface, and therefore asymmetric with respect to the midfrontal plane. Because fully congruent meniscal bearings have shown a very small wear rate at the knee,37,41 a thick PE meniscus is unnecessary, so that the overall thickness of the prosthesis, and therefore of the total amount of bone to be removed, can be minimized. The thickness of the central part of the component varies in 1-mm steps from 5–8 mm.
The three components are available in three different sizes. These were determined based on unpublished data obtained by us from measurements on lateral radiographs taken from numerous consenting volunteers, on previous reports of talocrural joint morphologic features,13,21 and on currently available prostheses.
The shapes of the three components allow unresisted relative motion by sliding on each other without separation or interpenetration while the isometric ligament fibers rotate about their origins and insertions without stretching or slackening. Particularly, to accomplish the necessary sliding and rolling guided by the isometric ankle ligaments, the meniscal-bearing element slides backward on the tibial component during plantar flexion and slides forward during dorsiflexion. The complementary nature of the engaged surfaces of the components is such that a large contact area is achieved in all positions. For the first time in total ankle arthroplasty, the multiaxial pattern of movement of the natural joint is replicated without significant distortion of the natural controlling and stabilizing mechanism while, at the same time, large contact areas are maintained throughout the motion arc. Because of the convex shape of the tibial component and because of the necessary compatibility of the articular surfaces with isometric ligament fiber rotations, this natural movement is obtained with an arc of curvature of the talar component significantly different from that of the natural anatomic shaping, and therefore from that of all the currently available meniscal-bearing designs.
A convex shape of the tibial component improves the level of AP entrapment of the biconcave meniscal bearing, with lower and upper independent and cumulative levels of entrapment for this freely mobile component. Mediolateral entrapment of the meniscal bearing is guaranteed also by the sulcus running on the dome of the talar component, avoiding the sharpened limiting interfaces as used in some of the previous designs to prevent dislocation and separation.
The Surgical Technique
For this new prosthesis to work properly, the three components must be implanted in an exact absolute position with respect to the ligament geometry. Particularly, for the meniscal bearing to slide smoothly on both the components, it is necessary that a constant gap is maintained between the tibial and the talar components over the motion arc.
A vertical anterior incision is made. The horizontal surface on the top of the talus first is prepared by removing a predetermined thickness of bone using a tibial alignment jig and mounting a talar cutting block to guide the saw blade. The level of tibial bone resection is determined by looking at the necessary minimum overall thickness of the prosthesis. After the flat tibial surface is prepared using a tibial cutting block and the holes for the two parallel bars are drilled, the most appropriate AP position for the talar component is determined by a step-by-step procedure using a talar template and a series of gap-gauges. In every tentative AP position, the gap between the tibial flat surface and the top of the talar template is measured in maximal plantar flexion and maximal dorsiflexion, using the gap-gauges. This procedure is repeated after removal of bone on the front of the talar dome, until these gaps are equal to within 1 mm. The talar dome then is prepared, with the anterior and posterior chamfers and the holes for the two pegs, all guided by the talar template. With the trial talar component implanted, the same step-by-step procedure is repeated for the determination of the most appropriate AP position of the tibial component. Final talar and final tibial components are implanted and the most appropriate meniscal-bearing component finally is inserted between them.
In Vitro Trial Implantation of Prototypes
Preliminary observations of the expected restoration of the original mechanism at the replaced joint were made as a pilot study. A comprehensive experimental in vitro validation of the new prosthesis currently is being done. Prototypes for the two metal bone-anchored components and for the PE meniscal bearings were manufactured. Surgical instruments for the proper preparation of the tibial and talar bones were devised and manufactured.
Trial implantation with these prototypes was done in six specimens from patients who had had below-knee amputation. During manually induced flexion of these specimens, the meniscal bearing was observed to slide backward on the tibial component during plantar flexion and forward during dorsiflexion, for a total run of approximately 5–7 mm (Fig 4). The sliding motion of the meniscal bearing over the tibial and talar components was smooth and continuous. Complete congruence among the three components was maintained over the entire motion arc, and also during manually induced internal and external rotation and inversion and eversion. The observations from trial implantation of prosthesis prototypes qualitatively validate the assumptions underlying the computer models of the intact and replaced ankle.
In Vivo Implantation in Patients
Prostheses recently were implanted in two female patients with posttraumatic osteoarthritis. The meniscal-bearing components were instrumented with three, 0.8-mm diameter tantalum beads, two in the front and one on the back surfaces, stuck on small PE pegs. The study was approved by the Ethics Committee of the Istituti Ortopedici Rizzoli. The patients were instructed about the new prosthesis and signed an informed consent. The new instrumentation worked well, and allowed accurate implantation of the three components in the target location. After implantation of the final components, the meniscal bearing was observed to move significantly anteriorly during dorsiflexion and posteriorly during plantar flexion over both the metal components. The motion was smooth and continuous, and complete congruence among the three components was maintained in all joint positions. Radiographs at maximal dorsiflexion and maximal plantar flexion and fluoroscanning confirmed the meniscal-bearing component moves in the direction and approximately for the distance predicted by the computer-based models. The patients are in their first month followup, and they will be assessed periodically by radiography and fluoroscanning.
The design of total ankle arthroplasty is a delicate balance between several contrasting criteria. Apparently, the essential compatibility between the retained ligamentous apparatus and the prosthetic articulating surfaces has not been addressed previously. A new total ankle prosthesis is proposed here to have articular surface shapes compatible with the isometric rotation of certain ligament fibers, as seen in the natural joint during passive flexion. Our fully congruent mobile-bearing prosthesis permits three-dimensional unresisted relative motion of the replaced ankle while maintaining full contact at the tibial-meniscal and meniscal-talar articulations in all joint positions.
The new design is not without limitations. Similar to currently available mobile-bearing ankle implants, our design is subject to instability in ligament-deficient joints. Although the risk of double-sided wear has been disproved in retrieval studies of meniscal knee replacements,37,41 the risk of dislocation of the meniscal bearing is a potential problem. In this respect, fully congruent and ligament-compatible articular surfaces may confer advantage over some currently available prostheses.
The AP length of the meniscal bearing is slightly shorter than that of any predecessors, with corresponding smaller contact areas at the component interfaces. The length was limited by the risk of impingement or overlapping for the predicted AP motion of the meniscal bearing. However, the convex shape of the tibial component has ensured an entrapment similar to the previous designs and possible lengthening can be considered after observations in replaced ankles.
The present convex tibial component may experience transmission of shear forces between the tibial and meniscal components, with an associated higher risk of tibial component loosening. However, in the currently available implants with a corresponding flat-to-flat interface, the resulting shear forces only are resisted by the ligaments, which must be well balanced. The appropriate ligament tensioning addressed in our design can ensure that the ligaments play their full role in transmitting the shear forces and may help to reduce bone-component interface shear forces. However, this theory can be proved only in future modeling and experimental work. Moreover, there may be concern about the thickness of the tibial component with the necessary larger tibial bone removal. The minimum overall thickness of our prosthesis (11–12 mm over the three sizes) is, however, in the range of the current three-component5,24 and two-component42,48 implants.
The design of the new prosthesis is based on our prior modeling investigations limited to the sagittal plane and to passive joint motion. Most of the novel features of the new design are, however, in the sagittal plane and motion is not restricted in the other two anatomic planes. Moreover, three-dimensional extension of the original two-dimensional models are being developed, and these are supporting all the previous two-dimensionally based findings. Because the function of the ligaments in controlling the passive mobility of the joint is restored, their physiologic contribution to joint stability also should be restored, as shown for the anterior drawer test.8 Additional experimental investigation, however, is needed. In addition, we will study the patients using gait analysis, radiostereometry, and fluoroscopy to ensure that restoration of natural mobility is matched by restoration of natural stability.
The fundamental difference between this design and its predecessors is that no constraints were imposed here to reproduce exactly the anatomic shapes of either natural articular surface. Currently available three-component prostheses5,24 aim at reproducing the shape of the talar dome but use a flat tibial surface to articulate with the meniscus. We have shown that those designs are unlikely to restore the characteristic original pattern of ligament tensioning.26,27
Values for the radii of the talar cup compatible with flat and convex tibial shapes are similar, and both are noticeably larger than natural talus and therefore larger than all predecessors. A convex tibial component was preferred in our prosthesis because it greatly increases the level of entrapment of the meniscal bearing, accommodating for the smaller entrapment provided by the larger radius of curvature at the meniscal-talar component interface. These articular surfaces can reproduce the natural pattern of relative motion of the corresponding segments even though the shapes of the natural tibial and talar surfaces are not closely reproduced as in most of the previous designs.
Differing from all previous three-component designs, our design allows for inversion and eversion in addition to internal and external rotation at the tibial-meniscal interface. This particularly is important when considering that the device should restore the characteristic motion of the entire ankle complex, comprising the ankle and the subtalar. Inversion and eversion do not require lift-off.
A new ankle prosthesis was designed to be compatible with the physiologic function of the retained ligamentous structures and to replicate original joint mobility. The device combines complete freedom from restraint with complete congruity of the components throughout the three-dimensional motion arc. Mathematical modeling of the joint has proved to be fundamental for this development. Motion exhibited in trial implantation agreed remarkably with that predicted from the modeling study. Our work has shown the fundamental role played by the ligaments in guiding joint motion in intact and replaced joints. Problems of previous designs might be related to incompatibility with the ligamentous structures.
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