Design of a Free-Floating Polycarbonate-Urethane

Design of a Free-Floating Polycarbonate-Urethane Meniscal Implant Using Finite Element Modeling and Experimental Validation Jonathan J. Elsner Sigal Portnoy Gal Zur Research and Development Center, Active Implants Corporation, Netanya 42505, Israel

Farshid Guilak Duke University Medical Center, Durham, NC 27710

Avi Shterling Eran Linder-Ganz1 e-mail: [email protected] Research and Development Center, Active Implants Corporation, Netanya 42505, Israel The development of a synthetic meniscal implant that does not require surgical attachment but still provides the biomechanical function necessary for joint preservation would have important advantages. We present a computational-experimental approach for the design optimization of a free-floating polycarbonateurethane (PCU) meniscal implant. Validated 3D finite element (FE) models of the knee and PCU-based implant were analyzed under physiological loads. The model was validated by comparing calculated pressures, determined from FE analysis to tibial plateau contact pressures measured in a cadaveric knee in vitro. Several models of the implant, some including embedded reinforcement fibers, were tested. An optimal implant configuration was then selected based on the ability to restore pressure distribution in the knee, manufacturability, and long-term safety. The optimal implant design entailed a PCU meniscus embedded with circumferential reinforcement made of polyethylene fibers. This selected design can be manufactured in various sizes, without risking its integrity under joint loads. Importantly, it produces an optimal pressure distribution, similar in shape and values to that of natural meniscus. We have shown that a fiber-reinforced, freefloating PCU meniscal implant can redistribute joint loads in a similar pattern to natural meniscus, without risking the integrity of the implant materials. 关DOI: 10.1115/1.4001892兴 Keywords: knee injury, prosthesis, computational model, contact pressure, cadaver, meniscectomy

1

Introduction

The menisci are semilunar wedge-shaped structures that play critical roles in the load distribution, shock absorption, and joint 1 Corresponding author. Contributed by the Bioengineering Division of ASME for publication in the JOURNAL OF BIOMECHANICAL ENGINEERING. Manuscript received April 21, 2010; final manuscript received May 24, 2010; accepted manuscript posted May 31, 2010; published online August 17, 2010. Editor: Michael Sacks.

Journal of Biomechanical Engineering

congruity in the knee 关1–3兴. Meniscal tears are common knee injuries that subsequently lead to degenerative arthritis, attributed primarily to the changes in the magnitude and pattern of stress distribution in the knee 关4–6兴. In such cases there is clearly a need to protect the articular cartilage by either repairing or replacing the menisci. Traditionally, meniscal replacement involves implantation of an allograft 关7兴. However, besides problems related to the availability, size matching, cost and risk of disease transmission, allograft menisci undergo remodeling after implantation, causing shrinkage and reduced mechanical strength 关8–10兴. These changes may lead to tears and dysfunction of the allograft, and thus contribute to uneven distribution of load, instability, and recurrence of degenerative damage 关11兴. A synthetic functional meniscal implant provides an appealing alternative that could have significant advantages for meniscal replacement. Specifically, an artificial meniscal implant that is designed to mimic the function of the natural meniscus 共e.g., replicate the normal stress distribution in the joint兲 could be available at the time of surgery in substantial number of sizes and shapes to accommodate most patients, while addressing other drawbacks associated with the use of allografts as mentioned above. The meniscus possesses highly complex structural, geometric, and mechanical properties that provide for its unique function. The natural meniscus contains collagen fibers, arranged predominantly in the circumferential direction, within a hydrated matrix 关39兴. This fiber arrangement supports the large hoop stresses that optimize distribution of contact stresses within the knee joint and prevent meniscal extrusion. An undesired extrusion of the meniscus, due to disruption of collagen fibers, may lead to knee pain and dysfunction by altering meniscal mechanics 关12兴. Therefore, the inhomogeneous and anisotropic properties of the meniscus play a critical role in its function 关13兴. The importance of the meniscal material properties was confirmed recently in a study by Haut Donahue et al. 关14兴, which showed, using finite element 共FE兲 simulations, that a meniscus implant with isotropic material properties could not adequately restore normal contact mechanics in the joint 关14兴. They concluded that transversely isotropic material properties, for example, as provided by a composite material, are required. A similar conclusion was reported by Vaziri et al. 关15兴, who suggested incorporating biocompatible fibers as an additional constituent material able to enhance circumferential stiffness of the meniscal implant and prevent meniscal extrusion 关15兴. This concept has in fact already been implemented in the design of several biodegradable scaffolds for meniscal regeneration. Kelly et al. 关16兴 designed a hydrogel construct reinforced with sutures that were woven through the implant. Chiari et al. 关17兴 designed an implant consisting of a poly 共␧-caprolactone兲 and HYAFF® blend, reinforced internally with poly-共lactic acid兲 fibers. Although both implants showed decreased cartilage degeneration in the short term in a sheep model, the former lacked essential durability and was torn radially as a result of failure of the matrix component, and the latter underwent compression and extrusion. In another study of a biodegradable polyurethane implant reinforced with carbon fibers in a canine model 关18兴, it was reported that carbon particles released from the implant induced a synovial inflammation, causing extra damage to the joint. These demonstrate the importance of material selection, and the major disadvantage associated with the meniscal tissue engineering concept: controlling the load transfer from the degrading scaffold material to the newly formed tissue while maintaining chondroprotection. The variability in the body response to biodegradable implants and the quality of the tissue formed still pose a problem in this respect, thus making it difficult to attain satisfying results from scaffold-type meniscal implants under intense knee loading conditions 关16,19–21兴. Furthermore, the proper attachment of an artificial meniscal implant may introduce a number of complexities with respect to both the design of the implant attachments as well as the surgical procedure. The development of a meniscal implant that does not require surgical attachment but still provides the

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biomechanical function necessary for joint preservation would have important advantages from several standpoints. Polycarbonate urethane 共PCU兲 is a tough, biologically stable polymer with a low elastic modulus 共10–100 MPa兲关22兴, in the same range as the dynamic properties of natural cartilage. It is an attractive material for a meniscal implant since it combines elasticity with durability and also offers good tribological properties comparable to those of cartilage 关22–24兴. Consequently, it has been in the focus of interest in recent years as an alternative to polyethylene 共PE兲 in orthopaedic load bearing applications in the hip 关22,24兴. As the primary matrix material of a meniscal implant, PCU has important advantages by providing the necessary mechanical properties and wear characteristics needed for long-term performance. Furthermore, a PCU implant can potentially be tailored to display more complex mechanical properties that replicate those of the native meniscus, such as transverse isotropy, which can be achieved by as a composite material through circumferential reinforcement. That is, the PCU itself would provide the load distribution and tribological function of the meniscal tissue matrix, whereas the circumferential fibrous reinforcement would take on the role of the circumferential collagen fibers that reinforce the natural meniscus, and thus, maintain internal hoop stresses, limit radial expansion, and contribute to improved pressure distribution. Therefore, the objective of this study was to design and evaluate different configurations of composite PCU-based medial meniscal implant. For this purpose, different reinforcement materials and configurations were investigated as parameters that may affect implant performance in terms of manufacturability, load distribution, implant extrusion, and long-term safety. Numerical simulations were performed to reduce manufacturing costs by negating mechanically problematic configurations and focusing on configurations that perform better. Analyses were performed by means of a validated FE model of the knee joint and PCU-based implant including circumferential reinforcement fibers under physiological loads.

2

Methods

An accurate cadaveric-based, three-dimensional 共3D兲 FE model was developed, in order to evaluate the biomechanical performance of different designs of an artificial self-centering, nonfixed, medial meniscal implant. The meniscal implant aims to replace a totally removed meniscus and thus was modeled within the joint space of totally meniscectomized medial compartment of the knee. Several models of the implant based on a PCU matrix, including embedded reinforcement fibers, were tested. Finally, a single implant configuration was selected, based on the implant’s ability to restore pressure distribution in the knee compared with natural intact meniscus and other restrictions such as manufacturability and long-term safety 共in respect to material loads兲. A schematic description of the aforementioned implant design and evaluation process is depicted in Fig. 1. 2.1 Finite Element Modeling. A 3D FE model of a representative medial compartment of the human knee was developed. The FE model geometry was reconstructed based on sagittal images of the knee region taken from magnetic resonance imaging 共MRI兲 scans 共GE Medical, 1.5 Tesla, T1 sequence, slice-thickness of 2 mm, gap of 0 mm兲 of a cadaveric left knee specimen 共male, 81 years, 82 kg; Figs. 2共a兲 and 2共b兲兲. The outer surfaces of the femoral 共Fig. 2共a兲兲 and tibial 共Fig. 2共b兲兲 articular cartilages were detected on each image and used to form the cross-sectional geometry. Next, cross sections were transformed to a 3D solid model by interpolation between MRI slices through the knee 共SOLIDWORKS; Fig. 2共c兲兲. Articular cartilage thickness was taken as 1.5 mm 关25兴, and the cartilage was assumed to be attached to the bone surfaces. The geometry of the FE-modeled implant was built based on the actual implant design 共NUsurface®, AIC, TN; Fig. 3兲, and the reinforcement fibers were modeled as “ring-shape” parts 共Fig. 095001-2 / Vol. 132, SEPTEMBER 2010

Fig. 1 Schematic descriptions of the medial meniscal implant design and evaluation process

2共c兲兲. Finally, the 3D models of the knee joint and implant were transferred to a FE solver 共ABAQUS 6.8, SIMULIA, RI兲 for nonlinear strain/stress analyses under loading conditions 共Fig. 2共d兲兲. The femur and tibia were assumed to be rigid surfaces. The articular cartilage and reinforcement fibers were considered as linear-elastic materials with elastic moduli and Poisson’s ratios, as detailed in Table 1. The main meniscal implant matrix, which is composed of PCU, was assumed to be incompressible hyperelastic material, and modeled using the Mooney–Rivlin material law, with the following energy function 关26兴: W = C10共I1 − 3兲 + C01共I2 − 3兲

共1兲

where the invariants of the principal stretch ratios ␭i are ¯I1 = ␭21 −2 −2 + ␭22 + ␭23 and ¯I2 = ␭−2 1 + ␭2 + ␭3 . C10 and C01 were taken as 2.912 and ⫺1.025, respectively, based on tensile, compression, and biaxial mechanical tests 共RAPRA Technology Ltd., UK兲. All nodes on the distal surface of the tibia were fixed for all translations and rotations 共Fig. 2共d兲兲. Based on the findings that peak knee loads during gait can amount to three times the bodyweight 关27,28兴, and that approximately 60% of this load falls on the medial compartment 关29兴, a 1200 N compressive load was used, as in the previous studies by Pena et al. 关30兴 and Zielinska and Donahue 关28兴. This vertical load was applied to the femur 共Fig. 2共d兲兲 at 0 deg flexion, the angle in which maximal load is applied to the knee during gait 关28,31兴. The meniscal implant was located as a free-floating device between the femur and tibia. In order to assure convergence of the numerical solution, the outer perimeter of the implant was fixed by a dozen springs with a very low stiffness modulus 共k = 0.05 N / m兲, all of which were connected to the ground 共Fig. 2共d兲兲. The influence of the springs on stresses developed on the outer shell of the implant was found to be negligible. We assumed an ideal adhesion between the reinforcement fibers and the PCU matrix. A frictionless contact condition was defined between the articular cartilages, and the superior and inferior surfaces of the implant. Penetration of nodes into another surface, which hinders tensile stress transfer across the interface, was prevented by using a contact pressure-clearance relationship defined as a “hard” contact model 关32兴. Transactions of the ASME


Fig. 2 The geometries of the medial meniscus and the „a… femoral and „b… tibial surfaces contacting it were extracted from sagittal MRI scans of a cadaveric male left knee to produce a „c… three-dimensional solid model of the medial knee. The solid model was then exported to „d… a finite element solver where it was loaded with 1200 N. All nodes on the distal surface of the tibia were fixed for all translations and rotations. The outer perimeter of the implant was fixed by a dozen springs „stiffness modulus k = 0.05 N / m… connected to the ground.

Rigid surfaces representing the bones were meshed with fournode quadrilateral surface elements 共3056 elements兲. The cartilage tissue and the implant were meshed with 3D hexahedron eightnode elements 共6112 and 13029 elements, respectively兲. The reinforcement fibers were meshed with two-nodes linear 3D truss elements that can bear only tensile stresses. Several configurations of the implant were tested for optimization purposes. First, a model of the meniscal implant with no reinforcement fibers was analyzed. Subsequently, reinforcement fibers were added, and the first factor considered was the material used as reinforcement. Six materials were evaluated: PE, carbon, Kevlar™, nitinol 共nickel titanium alloy兲, titanium, and stainless steel 316 共Table 1兲. The platform for this test was an implant with three fibers loops, each containing fibers with a total crosssectional area of 0.314 mm2. Next, the effect of the distribution pattern of the reinforcement material on the loads developed in the PCU implant was evaluated by the modification of the number of

loops containing the fibers. A total of 21 fibers of the same material were distributed between one, three, or seven loops in these analyses 共Table 2兲. Finally, the effect of changes in the total fiber volume incorporated in the PCU matrix on the loads developed in the implant and fibers was evaluated in two additional models containing a total of 30 and 39 fibers per cross section 共Table 2兲. For each configuration, tibial plateau 共TP兲 contact pressure distributions, TP contact areas, circumferential expansion of the PCU, and internal peak stresses were calculated. Plots of the contact pressure maps and internal stress distributions were produced for each simulation case. In addition, internal average and peak strains/stresses developed in the PCU and in the reinforcement fibers were calculated during loading. Then, a final optimal configuration of the implant and its reinforcement fibers were selected. The criteria for the selection of the final implant configuration were as follows: 共i兲 design for manufacturability in all sizes

Table 1 Material properties of the bones, articular cartilage, and different fibers Elastic Poisson’s modulus ratio Reference

Model component

Fig. 3 The meniscal implant composed of a polycarbonateurethane „PCU… matrix, reinforced with circumferential polyethylene „PE… fibers

Journal of Biomechanical Engineering

Bones 共femur and tibia兲a Articular cartilage Polyethylene fibers 共diameter= 0.2 mm兲 Carbon fibers 共diameter= 0.2 mm兲 Kevlar fibers 共diameter= 0.2 mm兲 Nitinol fibers 共diameter= 0.127 mm兲 Titanium fibers 共diameter= 0.127 mm兲 Stainless steel 316 fibers 共diameter= 0.127 mm兲 a

Rigid 15 MPa 98 GPa 234 GPa 112 GPa 83 GPa 116 GPa

Rigid 0.475 0.36 0.36 0.36 0.3 0.34

关38兴关40兴关35兴关41兴关41兴关42兴关42兴

179 GPa

0.3

关42兴

Bones can be considered as fully rigid with respect to the relatively soft cartilage.

SEPTEMBER 2010, Vol. 132 / 095001-3


Table 2 Peak and average strains and stresses in the finite element analyses, simulating reinforcement loop numbers of 0, 1, 3, 5, and 7 with different numbers of fibers per loop, under 1200 N compression PCU Parameter Simulation cases Total Fibers fibers disperse 21 21 21 30 39 a

No fibersa 21 7,7,7 3,3,3,3,3,3,3 10,10,10 6,10,10,9,4

Fibers

von Mises stress 共MPa兲

Compression strain 共%兲

Tensile strain 共%兲

Compression stress 共MPa兲

Tensile stress 共MPa兲

Peak

Avg.

Peak

Avg.

Peak

Avg.

Peak

Avg.

Peak Avg.

Peak

Avg.

Peak Avg.

Peak

Avg.

10.5 8 7.2 9.1 7.7 7.6

1.4 1.3 1.5 1.5 1.5 1.4

3.6 2 4.6 3.1 4.6 3

1.2 0.3 0.2 0.3 0.4 0.4

8.0 4.3 4.1 4.7 4.9 4.5

2.4 1.3 1.2 1.3 1.4 1.4

92 46.8 45.4 46.1 44.6 45.2

24.5 9.7 9.8 10.2 10.6 10.5

65.7 29.6 28.9 30.9 30.7 30.4

282.1 475.9 755.9 439 869.1

227.1 255.8 255.9 171 179.3

0.29 0.49 0.57 0.38 0.89

4.38 6.27 6.89 6.72 6.45 6.48

1.17 0.99 1.0 1.0 0.99 1.0

18.9 8.9 8.1 9.3 9.9 9.8

Tensile stress 共MPa兲

TP

Tensile strain 共%兲

0.23 0.26 0.34 0.18 0.42

Contact pressure 共MPa兲

Due to excessive deformations in this model, the analysis could not be completed 共up to 1200 N兲 so data were extrapolated from a loading of 1000 N.

共ii兲, loads in the PCU and fibers were within the manufacturer’s allowed limits, and 共iii兲 the designed meniscal implant model produced an optimal pressure distribution, compared with the natural meniscus under compression. 2.2 In Vitro Validation. The FE model of the final implant configuration 共Fig. 3兲 was validated by comparing the TP contact pressures measured in a cadaveric knee, in vitro, to the calculated TP contact pressures, which resulted from the FE analysis. The same knee used for the model development was fixed with bone cement to specially designed holders in a knee compression apparatus 共Tinius Olsen, Inc., PA; Fig. 4共a兲兲. The natural knee alignment was maintained under 0 deg flexion, and contact pressures under the intact meniscus were measured using thin and flexible contact pressure sensors 共K-scan™, Tekscan Inc., Boston, MA; Figs. 4共a兲 and 4共b兲兲 under loading conditions similar to those

Fig. 4 „a… The setup for the in vitro experiments. The femur and tibia of a normal human cadaveric knee was fixed with bone cement to specially designed holders in a knee compression apparatus. Contact pressures under the intact medial meniscus or the implant were measured using „b… force sensors while subjecting the knee to 1200 N vertical compression.

095001-4 / Vol. 132, SEPTEMBER 2010

described for the FE model. The chosen implant configuration, manufactured for this purpose, was located in the joint space, and contact pressures under the implant were measured. A peripheral path 共Fig. 4共b兲兲 was defined on the pressure map of both configurations 共measured and simulated兲, and pressures along this path were presented. We then compared pressure patterns along paths I 共measured兲 and II 共calculated兲. Statistical agreement between the both was evaluated by calculating the Pearson correlation coefficient 共r兲关33兴. A more detailed description of the experimental setup and procedure has been described previously 关34兴.

3

Results

3.1 Optimization of the Composite Material Parameters. The compression of an initial implant design, composed solely of PCU, resulted in relatively large distortions in the implant structure and peripheral extrusion with respect to reinforced implant configurations. Predominantly, the circumferential expansion of the nonreinforced implant under load was nine-times greater 共17.2 mm兲, compared with reinforced configurations. In fact, due to excessive deformations in this model, the analysis could not be completed up to 1200 N so that data were extrapolated from a loading of 1000 N 共Table 2兲. Internal peak compressive and tensile strains in the PCU were found to reach 92% and 67%, respectively, and the average von Mises stresses within the PCU material reached 8 MPa. These values are at least two times higher than those calculated in the rest of our simulations 共Table 2兲. Interestingly, although the average TP pressure was higher in the nonreinforced configuration, peak TP pressure was lower than in the reinforced PCU 共Table 2, Figs. 6共a兲 and 6共c兲兲. Thus, subsequent FE simulations of the implant model included circumferential reinforcement with various biocompatible fibrous-form materials, as listed in Table 1. The results for PCU embedded with different fibrous reinforcement materials did not, however, differ in terms of the mechanical conditions in the implant or its pressure distribution capability on the TP. Specifically, the peak and average tensile strains in the reinforcement fibers were 0.36⫾ 0.1% and 0.14⫾ 0.05%, respectively. The peak and average compression stresses in the PCU for the various reinforcements were 7.7⫾ 0.1 MPa and 1.5⫾ 0.1 MPa, respectively. Also, the peak 共6.5⫾ 0.1 MPa兲 and average 共1 MPa兲 TP contact pressures were practically identical for all simulated materials. Consequently, the next simulations were focused on implant structures reinforced with PE fibers 共Dyneema Purity®, DSM兲关35兴. These fibers are presently used for medical applications, e.g., orthopaedic augmentation and suturing, owing to their excellent wear resistance and long-term mechanical stability. The effect of reinforcement 共PE fibers兲 dispersal in the matrix 共PCU兲 was evaluated using the FE model by modification of the distribution of 21 PE fibers through the implant’s cross section. Transactions of the ASME


Table 3 Calculated loads in the selected implant configuration versus allowed values Measure Cartilage pressure 共MPa兲 PCU compression stress 共MPa兲 Fibers strain 共%兲 Fibers stress 共MPa兲

Calculated valuea

Allowed value

6.89 7.2 0.49 476

5–9b 15 共PTG corp.兲 3.4 共DSM corp.兲 3100 共DSM corp.兲

a

Based on the optimized implant configuration of three loops with seven fibers in each loop, compressed at 1200 N. b Based on measurements under intact menisci from eight cadaveric knees.

Similar strains and stresses were calculated in the main matrix component in all cases 共Table 2兲. Implant extrusion, as indicated by the increase in the external perimeter of the implant under load, was reduced by 50% when three and seven loops were used 共⬃0.93 mm兲, compared with the use of one loop only 共1.38 mm兲. Tensile strains/stresses in the fibers were found to be higher when more loops were simulated 共Table 2兲. However, all values were well below the manufacturer’s allowed limits 共Table 2 and 3兲. Based on these results and the production considerations 共discussed below兲, we focused our analyses on models containing three to five loops. Finally, we studied the effect of increased reinforcement volumes: from a total of 21 fibers dispersed evenly in three loops 共results presented in Table 2兲 to 30 共10 per loop兲 or 39 共six, ten, ten, nine, and four fibers per loop, starting at the proximal loop兲. These analyses yielded no apparent difference in both the peak and average strains and stresses in the PCU under compression. The circumferential expansion of the implant under load was mildly reduced to 0.91 mm in both new cases. The first model, based on an increase in the amount of fibers per loop 共from seven to ten兲, resulted in a 20% reduction in the peak tensile strains in the fibers. Similarly, the peak tensile stresses in the fibers were ⬃8% lower due to the increase in the fiber content. Unfortunately, we found that the integrity of the physical implant containing ten fibers per loop was compromised. The implant containing 39 fibers produced peak tensile strains and stresses in the fibers that were ⬃2-times higher than in the implant containing 21 fibers. The peak loads accumulated in the superior and inferior loops 共containing four and six fibers per loop, respectively兲. However, the average tensile stresses in the implant containing 39 fibers were ⬃30% lower than the average tensile stress in the 21 fiber configuration. The 39 fiber configuration was also at a disadvantage compared with the three-loop implant due to complications in the production process. Generally, the production success-rate for implants containing five loops or more than seven fibers per loop was poor, i.e., PCU adhesion was not strong enough to maintain integrity under load. Specifically, the number of disqualified implants with more than seven fibers per loop or implants with five or more loops was three times higher than the number of disqualified implants in a three-loop implant with seven fibers per loop 共in a nonmass production process兲. 3.2 Validation of the Computational Results. Contact pressures calculated from the FE analysis were similar to the measured TP contact pressures 共Figs. 5共a兲 and 5共b兲兲. Specifically, pressure concentrations were located in the same regions A, B and C, and regions with less contact were located around region D, in both cases. Pressure patterns along paths I and II were found to be analogous in terms of the peak values and locations 共Figs. 5共a兲 and 5共b兲兲. The Pearson correlation coefficient 共r兲 between the measured and calculated pressure-location curves, was 0.98 共p ⬍ 0.0001, Fig. 5共b兲兲. 3.3 Final Configuration Results. The implant containing three loops with seven PE fibers per loop was selected as the optimal implant design. The TP contact area for this implant was Journal of Biomechanical Engineering

approximately 712 mm2. This design fulfills all three restrictions mentioned before 共Fig. 1兲. Specifically, this design can be manufactured in various sizes, and the loads in the PCU and fibers satisfy the manufacturer’s limits, as shown in Table 3. Finally, this configuration produces an optimal pressure distribution 共Fig. 6共a兲兲, similar to that of natural meniscus 共Fig. 6共b兲兲.

4

Discussion

In this study, we presented an FE-based approach for design optimization of a nonfixated 共free-floating兲 meniscal implant based on PCU. The meniscus’ important role in the distribution of joint loads has long been recognized 关4–6兴. Consequently, the initial assessment of meniscal replacements, most often allografts, has been conducted by the measurement of load distribution patterns on the TP under quasi-static loading conditions representing peak gait loads. The data already available based on a combination of experimental 关14,29,36,37兴 and computational works 关14,15,32,38兴 have led to a good understanding of important biomechanical factors affecting the load distribution ability of the meniscus. Of these, it appears that the material properties and geometrical attributes, including size-matching to the knee, are the most important factors which ought to be addressed in the design of an artificial meniscal replacement. The advanced 3D MRIreconstructions of the geometry of natural meniscal and knee components, as demonstrated in this study, can provide adequate response to the latter issues when scheming the form and size range of a synthetic meniscal replacement. However, although attempts have been made to produce synthetic materials for the total meniscal replacement 关16–18兴, there is still no commercial implant available in clinical use that reproduces the load distribution ability of the natural meniscus and satisfies long-term requirements of stability and safety. PCU corresponds well to the basic requirements from a meniscal implant material, namely, similar mechanical properties to the dynamic properties of cartilage 共compressive modulus of 20 MPa兲 and durability to long-term loading conditions 关22兴. The ability of PCU to permit local material flow and consequently increase the contact area under joint loads is regarded as an advantage of this material compared with rigid bearing materials used in orthopaedics 共e.g., PE and metal兲. However, the present analyses of a meniscal implant composed solely of PCU exposed several limitations that needed to be addressed. First, this unreinforced meniscal implant underwent excessive deformation under compression 共circumferential expansion of 17.2 mm兲, which can potentially lead to adverse effects such as tibial overhang or impingement on the cruciate and collateral ligaments. Second, TP contact pressures under this implant were concentrated in the central region and did not retain a semilunar pattern characteristic of the natural meniscus 共Fig. 6共c兲兲. Evidently, material overflow beyond tibial margin should be avoided in order to effectively increase the contact area and better distribute the loads. These findings strengthen the previous theories regarding the need for anisotropic material properties, and in particular, increased circumferential stiffness, as in the natural meniscus, to restore meniscal function 关14,15兴. Thus, based on the results, it became clear that the selected configuration of a PCU meniscal implant must include reinforcement fibers to prevent undesired expansion. Importantly, circumferential reinforcement of the PCU implant, similar to the structural properties of the native meniscus, improved its performance. First, it considerably reduced circumferential expansion under load by ⬃95% to ⬃2 mm. Second, it resulted in an improved ability to distribute the TP pressure. The pressure distribution maps attained for reinforced PCU had a closer form to that attained for the natural meniscus. Importantly, the contact area was predominantly in the outer third of the TP surface and was not concentrated in the central region 共Fig. 6兲. Surprisingly, despite a diverse spectrum of elastic moduli 共83– 234 GPa; Table 1兲, the functional results of the various reinforceSEPTEMBER 2010, Vol. 132 / 095001-5


Fig. 5 „a… Contact pressure maps on the tibial plateau for calculated „left frame… and measured „right frame… configurations; „b… pressures along peripheral paths I „measured… and II „calculated… were plotted

ment materials chosen for the study 共Table 1兲 were similar, in terms of the internal loads 共matrix or reinforcement兲 and the implants pressure distribution ability. We hypothesize that the large disparity between the elastic moduli of the reinforcement materials and PCU matrix 共3–4 orders of magnitude; Table 1兲 overshadows milder differences between the properties of the reinforcement materials, thus leading to similar outcomes. Based on these findings we conclude that the choice of the optimal reinforcement material should be made according to other criteria, e.g., biocompatibility, long-term resistance of the reinforcement material to fatigue, adhesion of the reinforcement material, and matrix component and manufacturability. For example, the metal wires display the shape memory and can be tricky to wind during manufacturing. Polyethylene, which has a long history of use in biomedical applications, successfully meets all these requirements and was selected for the final implant production. Two fundamental parameters of the composite material, i.e., 共i兲 the distribution pattern of the reinforcing material in the PCU matrix, and 共ii兲 the volume of reinforcement incorporated, was examined as modulators of the overall mechanical environment within the implant and its ability to redistribute the joint loads 095001-6 / Vol. 132, SEPTEMBER 2010

over the TP. The distribution of fibers within an implant cross section had a mild effect on the stress-strain distribution within the composite material. We found that a single loop containing 21 fibers was least effective in converting compressive joint loads into tensile stresses acting on the fibers. This might be caused by a relatively small fiber to the matrix interface area through which forces can be mediated. The peak and average strains 共0.29% and 0.23%, respectively兲 and stresses 共282 MPa and 227 MPa, respectively兲 in the fibers in this configuration were the lowest 共Table 2兲. Dispersal of a similar amount of fibers 共21兲 throughout the crosssectional area 共in three or seven loops兲 increased the average tensile stress in the fibers 共Table 2兲. This finding indicates a more active role of the fibers in load conversion by sustaining the hoop stresses. Interestingly, the maximal spreading of fibers, set in seven loops, cannot be considered as useful since peak stresses in fibers located near the superior/inferior implant surfaces were extreme 共776 MPa, Table 2兲, which is not a desirable outcome. Additionally, there was no improvement in the load bearing function of the fibers 共average tensile stress remained 256 MPa兲. Hence, we concluded that the optimal number of loops should be between three and five. Transactions of the ASME


Fig. 6 „a… Calculated contact pressure maps under the selected reinforced meniscus implant loaded from 0 N to 1200 N, „b… corresponding contact pressure maps measured in vitro under an intact natural meniscus loaded from 0 N to 1200 N, „c… calculated contact pressure maps under a nonreinforced meniscus implant loaded from 0 N to 1000 N „due to excessive deformations in this model, the analysis could not be completed up to 1200 N…

Finally, we examined the effect of increased reinforcement volume in the PCU matrix. We found that an increase in the fiber content can be beneficial in reducing average fiber loads. For example, when increasing the amount of fibers from 21 to 30 共in a similar fiber distribution of three loops兲, the load per fiber was reduced in an inverse proportion to the increase in fiber numbers. The increase in the fiber volume did not, however, affect the total load borne by the fibers. On the other hand, increasing the amount of fibers to 39 required the usage of additional loops, which subsequently increased the peak stresses and strains in the fibers 共Table 2兲, as previously mentioned. Furthermore, from a practical standpoint, we found that an increase in the fiber content 共above 21兲, increased the risk for faults in implant production, and therefore, could not be implemented, especially in small implant sizes. A limitation of the FE models are the assumptions made regarding the mechanical properties, as discussed in Sec. 2. These assumptions, however, are reasonable since the scope of this study was to assess the effects of different implant configurations on the mechanical state of the medial knee. Additionally, both the FE analyses and the in vitro studies performed were quasi-static simulations. We believe that the examination of the meniscal performance at the most extreme phase of gait, as done in this study, can provide a measure of potential abnormalities in pressure distribution patterns. Further simulations of dynamic loading are advised for studying the behavior of the implant during activities such as gait, ascending and descending stairs, running, and jumping. In spite of the mentioned limitations, the free-floating implant design proposed herein has several advantages. We have shown that a reinforced PCU meniscal implant can redistribute the joint Journal of Biomechanical Engineering

loads in a similar pattern 共shape and values兲 to the natural meniscus, without risking the integrity of the implant materials. Importantly, the discoid-shaped implant design does not require fixation to the tibia, as do allografts, and thus, can be inserted using minimal incision with minimal complications. As a result, this procedure can be considered a bone-spearing solution and allows multiple possibilities for future treatment options. We conclude that our selected implant configuration can potentially serve as a meniscal implant in clinical settings.

Acknowledgment The authors wish to acknowledge the International Institute for Advancement of Medicine 共IIAM, Jessup, PA兲 for providing human tissue samples used in this study. The work was funded by Active Implants Corporation.

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Transactions of the ASME
