Total hip replacement (THR) is an effective treatment option to manage painful hip diseases in dogs, such as coxofemoral osteoarthritis secondary to hip dysplasia, traumatic coxofemoral luxation, avascular necrosis of the femoral head, and comminuted fractures of the femoral head.1–7 Complication rates after THR range from 5% to 30%.3,5,8–13 This variability can be explained to some extent by the different implant systems, the generation of implants, the surgical technique, the surgeon's experience, and the duration of follow-up of the studies.
Both cemented and cementless prostheses are commercially available. Cemented prostheses provide immediate postoperative stabilization of implants within the bone and lead to rapid patient recovery with the resolution of clinical signs. However, cemented prostheses are associated with a risk of late failure after aseptic loosening at the bone-cement or cement-implant interface.14–17 To avoid these complications, cementless prostheses were developed. The stability of cementless prostheses initially relies on implant impaction through the press-fit effect and then on long-term osseointegration. However, 5% to 33% of postoperative complications are related to the acetabular component of cementless prostheses,5,11–13,18,19 and these complications often require revision. A good press-fit and positioning are essential to prevent the risk of postoperative complications.
Primary stability of the press-fit system is achieved by impacting the implants into a prepared bed that is slightly smaller than the implants. Characteristics of the cup also play an important role in primary stability through a combination of friction effects, rim effects, and geometry (surface and shape). This initial stability is mandatory to achieve long-term osseointegration.
Several cup designs have been developed to improve press-fit quality in human medicine. Human studies have shown superior in vitro stability of cups with enhanced fixation provided by geometric designs with surface fins.20,21 However, some concerns about the required impaction force have been raised when large fins are involved.22–24 Currently, in human medicine, the hemiellipsoidal cup (hemispherical with a flattened dome) is the most commonly used geometrical design to achieve the press-fit effect, mainly because of the ease of acetabular bed preparation and reproducibility of insertion.25–27
In veterinary medicine, a cementless acetabular implant with reinforced fixation around the periphery would have smaller fin sizes because of the smaller implant size than those tested in human medicine. The density of human trabecular bone is quite similar to that of dogs.28,29 Moreover, canine trabecular bone has a lower modulus than the human bone of equal density and, therefore, can reach higher compressive stresses before breaking.30 Therefore, the addition of a peripheral fin might improve primary stability in dogs undergoing THR. Although several veterinary studies have described the stability of selected acetabular cups in different load and implantation conditions31,32, none have described the effect of cup geometry on primary stability. Currently, only slightly hemiellipsoidal acetabular implants have been developed and are available for dogs.
The purpose of our study was to compare the in vitro mechanical performance of cementless acetabular implants with peripherally reinforced fixation to that of a conventional hemiellipsoidal model and to determine the effect of immediate press-fit and resistance to applied forces by testing the limits of the implants under extreme conditions. The 2 implants were designed from a traditional hemiellipsoidal model with peripheral fixation surface fins. We chose to add peripheral fins to test enhanced primary stability due to a stronger grip on the periacetabular bone while allowing the porous surface of the cup to be flush with the underlying bone, with few bone-prosthesis gaps. The null hypothesis was that the surface geometry effect of acetabular implants would not affect the initial stability under extreme loading. The specific aim was to assess the initial stability using the edge-loading and push-out tests and then compare the implantation behavior of 3 acetabular cups with different surface geometries that affect edge effects.
Methods
Acetabular implants
We used 3 distinctive canine acetabular implant designs (Portevet Inc; Figure 1). The implants, which were designed and made specifically for this study, were based on a hemiellipsoidal design with part of the equatorial rim truncated at a 40° angle, 3.8 mm from the center of the radius. This portion was positioned on the dorsal edge to match the geometry of the canine acetabulum. The equatorial diameter of the implants was 28.5 mm and the polar radius was 13.3 mm. The implants were all made of a titanium alloy (Ti6Al4V) with a 150-µm ± 50-µm porous coating using plasma spray technology. The implants differed in their surface geometry. Model A was the standard reference model. Model B had peripheral fins creating a discontinuous bead with a 0.70-mm radius, positioned 2 mm from the rim of the implant. Model C had peripheral fins organized into 2 levels of discontinuous beading with a 0.70-mm radius, one located at the same level as in Model B and the second level located 3.5 mm below (Figure 1). The interior was made of a polyethylene liner that clips onto the metal back to accommodate a 16-mm diameter head.
Photograph of the 3 implant designs with, from left to right, Models A, B, and C
Citation: American Journal of Veterinary Research 84, 9; 10.2460/ajvr.23.03.0066
We had 3 implants of each model. Each implant was used up to 10 times. Cups were manually cleaned after each test to remove any debris that may have affected the texture of the cup surface. A visual inspection of the cup surface was also performed after each test to ensure that it was not scratched or damaged.
Synthetic bone substrate
The study was conducted using a polyurethane (PU) foam block that reproduced the density of canine trabecular bone and compression modulus. This model was developed based on previous in vitro testing of human21,33,34 and canine31 acetabular implants in PU foam blocks.
PU foam blocks with a density of 0.32 g/cm3 (20 pcf) and a compression modulus of 210 MPa were used as a substrate (Sawbones Inc). Each block was homogeneous and measured 65 X 60 X 40 mm with a truncated edge at a 40° angle, made to reproduce the dorsal shape of the canine acetabulum. Acetabulum-like beds were prepared manually by sequentially using a 16-mm diameter wood drill bit to center the cavity and then dedicated cutting and finishing reamers (Portevet Inc) to create a depth of 14 mm. Reaming was made coaxial with the axis used for the impaction of the cup. Each block was used once and randomly assigned to a specific test and model. For the push-out tests, a 16-mm diameter drilling hole centered on the acetabulum-like bed was added through the entire height of the substrate.
Implantation
For the first set of edge-loading tests, cups were manually impacted into PU blocks by a board-certified surgeon (TC) with the same instrumentation as in a clinical setting.
For the second series of edge-loading tests and push-out tests, cups were impacted in a standardized manner using a uniaxial material testing machine (AGS-X, Shimadzu Inc) to reduce the variability associated with the manual impaction method. The sensors used to record the force (10 kN load cell) and the displacement (mechanical traverse stroke) were those natively associated with the testing machine. Synchronized acquisition of the measurements was carried out using the Trapezium X software (Shimadzu Inc) with a sampling frequency of 10 Hz. Cups were manually positioned in the correct orientation before impaction. A custom-made machine-mounted 16-mm diameter femoral head was centered and positioned in the polyethylene inner layer of the cup. A 10-cycle quasi-static compression at 1200N was then performed based on the results of a previous ex vivo canine study.35
The cups were oriented and stabilized at a 45° “lateral opening” angle and 0° “version” and “inclination” angle. Acetabular components were considered to be in place when the entire metal back area of the component was engaged in the bone substrate, such that only the inner sleeve of the acetabular component was visible. For all specimens, only cups that were correctly placed with the equatorial rim of the metal matching the upper surface of the substrate and the truncated dorsal rim were included.
Test modes
Mechanical testing was carried out using the 10 kN uniaxial material testing machine, with the foam blocks secured in a clamp.
Edge-Loading Test—Eccentric edge loading was applied with a metal indenter (Figure 2). The tests were performed under a moving load at a speed of 1 mm/s until failure. Failure was defined as cup displacement or block fracture. A set of 18 trials was performed, in which 6 specimens of each acetabular cup design were manually impacted. A second series of 15 trials were performed with 5 specimens of each acetabular cup design impacted in a standardized manner.
(A) Edge-loading test setup, (B) push-out test setup. View of the backside of the PU foam block with the dome of the cup protruding.
Citation: American Journal of Veterinary Research 84, 9; 10.2460/ajvr.23.03.0066
Push-Out Test—The foam block with the impacted cup was placed upside down on a clamp that did not interfere with the displacement of the cup. A metallic indenter was centered on the dome of the cup and protruded from the back of the foam block through the hole (Figure 2). Compression at a rate of 1 mm/s was performed until the cup was dislodged from the cavity and the maximum failure load was recorded. Five foam blocks were used per group.
Static Seating Force and Implantation Behavior—The static seating force of the different acetabular cup designs was compared using the uniaxial material testing machine. Cups were manually positioned unimpacted to have the correct orientation in the cavity before impaction. A custom-made machine-mounted 16-mm diameter femoral head was centered and applied against the polyethylene liner of the implant until a preload of 10 N force was recorded. The insertion device was not fixed to the cup and allowed rotation of the cup. The height of the cup protruding from the surface of the PU foam block was approximately 3 mm. Force and displacement recording measurements were set to 0 for this initial positioning condition. A force-displacement curve recorded up to 3 mm of displacement of the stroke of the machine at a 2 mm/min speed. The impaction process and cup positioning were visually checked before and during the test. Fifteen trials were performed with 5 specimens per group.
Statistical analysis
The sample size was too small to prove normality, therefore, nonparametric tests were used. Differences in edge-loading and push-out forces among groups were assessed using the Kruskal-Wallis test. An alpha level of 0.05 was used for all tests of significance (significance at P < .05). The null hypothesis was that there was no effect of surface geometry according to the test performed (edge-loading or push-out). When the null hypothesis was rejected in the Kruskal Wallis test analysis, comparison tests based on confidence interval were conducted. All statistical analyses were carried out using XLSTAT (Addinsoft Inc).
Results
Edge-loading test with manual impaction
The mean load at failure of manually impacted cups was higher for Model A cups than for Models B and C. However, there were no statistically significant differences in maximum edge-loading force among the 3-cup designs (H(2) = 0.041, P = .980; Figure 3). The mode of rupture was a periprosthetic fracture of the PU foam block simultaneous with the displacement of the cup for 2 Model C specimens and 1 Model B specimen. The other cases had a 1-mm displacement. When the Model B and C cups were removed from their cavities, there was a circumferential depression in the substrate corresponding to the cup rim region.
Box plot comparing the differences in maximum edge-loading force (Fmax) between the 3 geometries (Models A, B, and C). (A) with manual impaction. (B) with machine impaction.
Citation: American Journal of Veterinary Research 84, 9; 10.2460/ajvr.23.03.0066
Five specimens had to be excluded before the trial, including 2 Model B and 2 Model C cups that tilted toward the dorsal edge during implantation, resulting in malpositioning and excessive opening of the lateral opening angle, and 1 Model C specimen that underwent periprosthetic block fracture during implantation. Therefore, a new impaction on new PU foam blocks had to be performed for 2 Model B implants and 3 Model C implants to have the right number of specimens per group for testing.
Edge-loading test with standardized impaction
There were no problems with the impaction of the acetabular cups. The variation in the model showed a statistically significant effect on the maximum force at failure (H(2) = 6.480, P = .0039; Figure 3). Interval confidence comparisons showed a statistically significant difference between design A and B . Model B showed a lower peak force than Model A with a peak force of 438.6 N 95% CI [412.8, 464.5] and 502.0 N 95% CI [477.4, 526.8], respectively.
For 3 Model C specimens, rupture consisted of a periprosthetic fracture of the PU foam block with a simultaneous displacement of the cup. A circumferential depression in the substrate was observed when Model B and C cups were removed after testing.
Push-out test
There were no problems with the impaction of the acetabular cups. The model design had a statistically significant effect on the maximum force at failure (H(2) = 9.420, P = .009). A significant difference was shown between the results of Model A (213.7 N 95% CI [193.2, 234.3]) and those of Models B and C (139.4 N 95% CI [130.2, 148.6]) and (138.9 N 95% CI [116.7, 161.2]), respectively; (Figure 4).
Box plot comparing the differences in maximum push-out force (Fmax) between the 3 geometries (Models A, B, and C).
Citation: American Journal of Veterinary Research 84, 9; 10.2460/ajvr.23.03.0066
The PU foam periprosthetic rim was systematically torn off during cup expulsion for Models B and C.
Static seating force and implantation behavior
There are 2 patterns in the force-displacement implantation curves, the first corresponds to Model A, and the second to Models B and C (Figure 5). There is a difference in stiffness between the 2 trends. For a 2-mm penetration, the mean required force was 194.4 ± 31.7 N for Model A, but approximately double for Model B (362.0 ± 42.6 N) and Model C (361.6 ± 27.5 N). Therefore, the addition of peripheral fins leads to a statistically significant difference in the static seating force required to impact the implant (H(2) = 9.420, P = .009), with significant differences between Model A (95% CI [166.6, 222.2]) and Models B and C (95% CI [324.6, 399.4] and 95% CI [337.5, 385.7], respectively).
Force (N) vs transverse stroke displacement (mm) curves for 5 Model A (solid line), 5 Model B (dashed line), and 5 Model C (dash-dotted line) implants.
Citation: American Journal of Veterinary Research 84, 9; 10.2460/ajvr.23.03.0066
Stiffness continued to increase for Model A, corresponding to the continuation of impaction and stiffness of the structure. In contrast, for Models B and C, the curves flattened, corresponding to a rotation of the cup in the foam block around the femoral head in compression. The rotation was systematically directed towards the dorsal edge and started after the equatorial peripheral fin had begun to engage the cavity. After cup removal, circumferential depression in the substrate was systematically observed for Models B and C, while no defects were seen for Model A cups (Figure 6).
Photograph of the circumferential depression (black arrows) in the substrate corresponding to the cup rim region after removal of a Model B cup following static seating test.
Citation: American Journal of Veterinary Research 84, 9; 10.2460/ajvr.23.03.0066
Discussion
This study has shown that peripheral fins do not improve the primary stability of acetabular cups during edge-loading and push-out tests, and may instead be deleterious by causing substrate damage in our synthetic bone model (Figure 6).
The edge-loading test with manually impacted acetabular cups did not show any effect of model geometry on primary stability. We decided to impact the cups manually for the first series of edge-loading tests as this replicates the insertion method used in surgery. As we were not able to guarantee an identical impact force for each specimen, it is possible that some of the variability between specimens is due to a slight variation in the impaction force used and the positioning of the cup. The exact force required to impact a cup into a correctly reamed acetabular cavity in vivo is currently unknown. The manual impaction force was determined to be approximately 1.2 kN in an ex vivo veterinary study.35 To reduce the variability associated with implantation, we decided to perform subsequent impactions using a machine with 10 compressive cycles of 1.2 kN. Similarly to our initial manually impacted edge-loading test results, the geometry of the acetabular cups had only a minimal effect on primary stability in the machine-impacted tests and the standard deviations were considerably reduced. These data show that the progressive insertion method provided a satisfactory and more repeatable seating than manual implantation, even though it is different from the method used in the clinic, in which a series of blows are applied. Regardless of the mode of impaction used, the added fins do not appear to affect the primary stability of properly impacted cups tested in edge-loading, unlike human studies with larger fins.20,21,36
During manual impaction, several specimens had to be excluded before testing due to poor orientation. The clinician noted more difficulty in impacting Model B or C implants than Model A implants. The implantation test did not replicate the impact insertion method used in surgery but allowed for a controlled insertion that provided an accurate measurement of the forces involved. We found that the force required for a given displacement was higher for the models with peripheral fins. Although the fins are smaller than those used in human medicine, their size is not negligible and these results are consistent with human studies.22,24,36–38 All Model B and C implants are tilted at the same level of insertion. This is probably secondary to the shape of the cup. The canine acetabulum is not completely hemispherical, so the cups were adapted with a truncated dorsal edge. Therefore, fins were only present on the cranial, ventral, and caudal edges. As the component is not axisymmetric, the distribution of frictional forces is higher where fins are present and the decrease in resistance at the dorsal edge may have caused the observed tilt. Although manual impaction by a series of strokes may reduce the risk of tilting in cups with fins, the force required is still likely to be greater than that required for a regular cup.
Push-out tests showed that peripheral geometry affects primary stability. Models with fins (B, C) showed a significant decrease in peak force associated with a periprosthetic fracture on expulsion. Peripheral fins damaged the substrate by cutting through it. We used only PU foam blocks to limit the variability between specimens, compared with cadaveric pelvises, allowing a more accurate comparison of the primary stability of different cup geometries. PU foam blocks have been widely used in previous studies.21,22,24,31,33,34,36–38 We selected a PU foam with a similar density and compressive modulus to that of cadaveric dog trabecular bone.28,29,39 Although the synthetic material was used as a representative model, it is unlikely that the substrate has the same viscoelastic response as real bone, despite having a similar compressive modulus. Furthermore, the foam blocks used in this study only model trabecular bone and do not replicate the cortical bone rim present at the periphery of the acetabulum. The subchondral bone envelope surrounding the cancellous region is important in the transmission of forces and fulcrum in the form of membrane stress,40 and therefore, the presence of a subchondral bone-like envelope could limit substrate damage. Moreover, the presence of a cortical shell may be associated with an even higher required impaction force for models with fins.
Implant designs with peripheral fins may increase the risk of poor impaction due to loss of press-fit, malpositioning, and periprosthetic fracture. Clinically, acetabular components should be implanted with accuracy and without excessive force. If high levels of force are necessary to impact the cup, this can lead to incomplete seating, malpositioning, and an increased risk of fractures. Periprosthetic acetabular fractures have been reported for uncemented implants in human medicine41,42 and 1 case has been reported in veterinary medicine.43 In addition, incomplete seating of the acetabular cup results in small residual compressive forces that can lead to excessive micromotion and impaired primary stability. Malpositioning may result in decreased range of motion,41,44,45 higher risk of luxation,46–48 and increased wear rates.49
We observed a circumferential depression matching the most equatorial fin after removing the cups during the push-out tests, which also suggests substrate damage during implantation. The most equatorial fins were placed almost flush with the edge of the acetabulum, but exceeded the diameter of the acetabulum, and, therefore, may have reduced the press-fit effect by damaging the substrate beyond its linear elasticity zone. Press-fit stability is a function of the contact area between the cup rim and the substrate.23 Design features may decrease total contact between the implant and the surrounding host bone. This reduces compressive forces and compromises stability, as has been shown for some implant geometries in human studies.33,36,37 This decrease in contact area also has consequences for long-term osseointegration, which is concentrated in pressure areas.50 Therefore, osseointegration may be focused only on the peripheral fins in Models B and C, instead of a larger contact area in Model A. We found no differences between Models B and C in our tests. This suggests that the second level of the fin in Model C did not cause substrate damage. The effect of this lower fin on press-fit could not be evaluated due to the presence of the first equatorial fin, which caused substrate damage.
To assess the fixation stability of each cup design, the limits of the implants were tested under extreme conditions, which simulate acute loosening with significant movement of the implant. The edge-loading and push-out tests assessed whether model geometry affected primary stability during the early postoperative period.21 These test modalities do not provide any information on micromovements, which should be assessed under physiological conditions, such as dynamic cyclic mode. Micromovements may induce the formation of a fibrous interface during a longer postoperative period, which can also lead to aseptic loosening.
Due to the lack of available implants, each acetabular component was re-used 10 times. Previous studies have shown that repeated use of implants up to 10 times did not affect the resulting primary impaction force and stability as long as the implants remained intact.31,33,35 We observed no obvious deformation or scratches on the implants and sequential analysis of the results did not show any differences in push-out force, edge-loading force, or static seating force in re-used implants.
Limitations of this study include those inherent to the use of a polyurethane model in which biomechanical analysis may differ from clinical results. Furthermore, we could only include a small number of trials per cup model due to the limited number of available implants. However, although a significant difference cannot be inferred, the results show that the peripheral fin does not improve primary stability under extreme loading and could be deleterious to the bone bed. The slight additional volume at the equatorial rim is sufficient to cause impaction difficulties by increasing the required forces and the risk of malpositioning.
To conclude, our results suggest an impairment of primary stability for acetabular cups with a peripheral design (B, C) compared with the hemiellipsoidal shape (A). Models with peripheral fin (B, C) appeared to cause incomplete seating if a higher force was not used during implantation and the risk of malpositioning was increased with this design. These results suggest that the hemiellipsoidal shape is superior to cups with peripheral fins, as it provides the same or better initial stability and requires a lower impaction force.
Acknowledgments
This work was supported by Portevet who provided the implants and instruments for implantation. However, Portevet was not involved with the experiments, data analysis, or reporting of this study.
The authors declare no conflicts of interest.
The authors thank Claude Carozzo, PhD, Dipl. ECVS, and Rolland Roume for their technical input and Michel Massenzio, PhD, for his advice on data interpretation. We thank Leah Cannon, PhD, for the English language editing of this manuscript. The support and encouragement of the entire Small Animal Surgical Department of the VetAgro Sup Small Animal Clinic, Marcy l'Etoile, France, are gratefully acknowledged.
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