Available online at www.sciencedirect.com

www.elsevier.com/locate/jmbbm

j o u r n a l o f t h e m e c h a n i c a l b e h a v i o r o f b i o m e d i c a l m a t e r i a l s 2 9 ( 2 0 1 4 ) 4 2 – 5 5

1751-6161/$ - see frohttp://dx.doi.org/10.

nCorresponding autE-mail address: jo

Viscoelastic properties of a syntheticmeniscus implant

Maoz Shemesha, Roy Ashera, Eyal Zylberberga, Farshid Guilakb,c,Eran Linder-Ganza, Jonathan J. Elsnera,n

aResearch and Development Center, Active Implants, 43 Hamelacha Street, Netanya 42505, IsraelbDepartment of Orthopaedic Surgery, Duke University Medical Center, Durham, NC 27710, USAcDepartment of Biomedical Engineering, Duke University Medical Center, Durham, NC 27708, USA

a r t i c l e i n f o

Article history:

Received 17 July 2013

Received in revised form

20 August 2013

Accepted 25 August 2013

Available online 3 September 2013

Keywords:

Polycarbonate-urethane

Prosthesis

Creep

Relaxation

Fatigue

nt matter & 2013 Elsevie1016/j.jmbbm.2013.08.021

hor. Tel.: þ972 9 8659220;n.elsner@activeimplants

a b s t r a c t

There are significant potential advantages for restoration of meniscal function using a bio-

stable synthetic implant that combines long-term durability with a dependable biomecha-

nical performance resembling that of the natural meniscus. A novel meniscus implant

made of a compliant polycarbonate-urethane matrix reinforced with high modulus

ultrahigh molecular weight polyethylene fibers was designed as a composite structure

that mimics the structural elements of the natural medial meniscus. The overall success of

such an implant is linked on its capability to replicate the stress distribution in the knee

over the long-term. As this function of the device is directly dependent on its mechanical

properties, changes to the material due to exposure to the joint environment and repeated

loading could have non-trivial influences on the viscoelastic properties of the implant.

Thus, the goal of this study was to measure and characterize the strain-rate response, as

well as the viscoelastic properties of the implant as measured by creep, stress relaxation,

and hysteresis after simulated use, by subjecting the implant to realistic joint loads up to 2

million cycles in a joint-like setting. The meniscus implant behaved as a non-linear

viscoelastic material. The implant underwent minimal plastic deformation after 2 million

fatigue loading cycles. Under low compressive loads, the implant was fairly flexible, and

able to deform relatively easily (E=120–200 kPa). However as the compressive load applied

on the implant was increased, the implant became stiffer (E=3.8–5.2 MPa), to resist

deformation. The meniscus implant appears well-matched to the viscoelastic properties

of the natural meniscus, and importantly, these properties were found to remain stable

and minimally affected by potentially degradative and loading conditions associated with

long-term use.

& 2013 Elsevier Ltd. All rights reserved.

r Ltd. All rights reserved.

fax: þ972 9 8659221..com (J.J. Elsner).

1. Introduction

The menisci of the knee are semi-lunar wedge-shapedcartilage structures whose main function is to increase jointcongruity between the femoral condyle and the tibial plateauand thus decrease stresses in the joint through the increasein the contact area (Kurosawa et al., 1980; Ahmed and Burke,1983; Levy et al., 1989; McDermott et al., 2008). Secondary

roles of the meniscus involve stabilization, lubrication, nutri-tion and proprioception (McDermott et al., 2008).

The meniscus is a biphasic composite material consistingof interstitial fluid (water and salts) and a porous and perme-able collagen and proteoglycan solid matrix. Water is themost abundant component, making up 63–75% of the totalweight of the meniscus (Herwig et al., 1984; Allen et al., 1995).Collagen is the principal constituent of the solid phase,

j o u r n a l o f t h e m e c h a n i c a l b e h a v i o r o f b i o m e d i c a l m a t e r i a l s 2 9 ( 2 0 1 4 ) 4 2 – 5 5 43

composing 75% of the dry weight (Allen et al., 1995). Thebiomechanical function of the native meniscus is closelylinked to its distinctive microstructure, and specifically tothe ability of circumferential bundles of collagen fibers,embedded within a hydrated matrix, to bear circumferentialhoop stresses. Biomechanical studies have determined thatthe collagen fiber arrangement substantially stiffens themeniscus in the fibers' direction, and that the inhomogeneityand anisotropy of meniscal structure results in significantregional and directional variations in stiffness and strength(Bullough et al., 1970; Fithian et al., 1990, Shirazi et al., 2008).The meniscus material is further complicated by the pre-sence of an appreciable viscoelastic mechanical responsebecause of interstitial fluid flow and intrinsic viscoelasticproperties of the tissue (Uezaki et al., 1979; Joshi et al., 1995;Chia and Hull, 2008). In such systems, when load impulsesare imparted to the tissue, energy dissipation can occur fromfluid movement in the tissue, which generates heat, orthrough rearranging of the molecular structure of the solidcomponents (Mak, 1986). Viscoelasticity is essential for theknee joint mechanical functioning under a wide range ofloading conditions, from low magnitude repetitive loads tohigh magnitude single impact loading, although the role ofthe meniscus as a shock absorber has been debated (Andrewset al., 2011). Cartilaginous tissue appears soft to facilitate jointmotion when not pressurized, but becomes highly pressur-ized under fast knee compression to support and redistributejoint loads (Ateshian, 2009; Kazemi et al., 2011). Experimentsshow that a combination of stress/strain rate, loading fre-quency, peak strain, and the number of total loading cyclesinfluences the mechanical behavior of cartilaginous tissues(Natoli and Athanasiou, 2009).

The instant mechanical response of the meniscus to load isgoverned by its tolerance to load and the severity of mechanicalimpact. A healthy meniscus is most likely to only undergo tearunder traumatic conditions (e.g., sports activities) (Terzidis et al.,2006). However the likelihood of a tear under non-traumaticconditions increases with degradation of the meniscus tissue asa result of aging (Englund et al., 2008) or prolonged exposure tooverloading as a result of obesity (Englund et al., 2008), occupa-tional loading (Baker et al., 2002; Rytter et al., 2009) or alteredbiomechanics (Lohmander et al., 2007). If sub-critical damagedoes not involve a significant tear to the meniscus tissue, themeniscus may attempt to repair or regenerate itself by synthe-sizing new extracellular matrix (ECM) components, or fixingdamage with material already present (Natoli and Athanasiou,2009). Either type of damage is likely to alter the biomechanicalperformance of the meniscus over time and compromise itsfunction (Natoli and Athanasiou, 2009; Katsuragawa et al., 2010;Bedi et al., 2010), and changes in the magnitude and pattern ofstress distribution in the knee as associated with a dysfunc-tional meniscus are associated with degenerative arthritis (Allenet al., 1984, McDermott and Amis, 2006).

Allograft implantation is one of the treatment options formeniscal replacement; however problems related to avail-ability, size matching, and cost, limit its use. In this regard,there are significant potential advantages to a bio-stablesynthetic meniscus implant that combines long-term dur-ability with a dependable biomechanical performance resem-bling that of the natural meniscus. The NUsurfaces Meniscus

Implant (Active Implants LTD, Netanya, Israel, Fig. 1a,b) wasdesigned as a composite structure that mimics the structuralelements of the natural medial meniscus. It is composed of acompliant polycarbonate-urethane (PCU, Bionates 80A, DSMBiomedical, Berkeley, CA) matrix which is reinforced circum-ferentially with high modulus ultrahigh molecular weightpolyethylene (UHMWPE) fibers (Dyneema Puritys, DSM Bio-medical, Geleen, NL). Laboratory and computational studieshave shown that the implant is functionally similar to themeniscus in terms of pressure distribution capability in itsvirgin state (Linder-Ganz et al., 2010 and Elsner et al., 2010a).

However, the overall success of such an implant is depen-dent on its long-term functional properties, and biomechanicaltesting of the device has not yet examined the effect of varying,time-dependent and dynamic testing conditions. In particular,changes to the material due to exposure to the joint environ-ment and repeated loading can have non-trivial influences onthe viscoelastic properties of the device (El Fray and Altstädt,2003; Geary et al., 2008; Kobayashi et al., 2005; Hannink et al.,2011). Thus, the aim of this study was to measure andcharacterize the strain-rate response, as well as the viscoelasticproperties of the implant as measured by creep, stress relaxa-tion, and hysteresis after simulated use, by subjecting theimplant to realistic joint loads.

2. Materials and methods

Two test models were used to simulate the long-term agingeffect of joint environment, static and dynamic. The firstmodel consisted of static (unloaded) soaking of the implant insimulated physiological fluid (SPF) which was prepared bydiluting bovine serum 1:4 to approximate protein concentra-tion of 1.6 g/100 ml with distilled water and maintained at37 1C. Fifteen implants were placed in SPF for 6 months.

The second model looked at the combined effect ofdynamic repeated loading and soaking in SPF for a totalduration of 2 million load cycles. Five implants were placed ina closed test cell containing SPF and loaded axially with acyclic load pattern that simulates gait (1500 N vertical load,applied at 2 Hz, for approximately 2 weeks). The load patternwas adapted from ISO 14243 (simulation of a total kneereplacement) by scaling down the force to 60%, to correspondto the relative load borne by the medial compartment (Winbyet al., 2009; Shelburne et al., 2006).

2.1. Fluid absorption measurements

Twenty NUsurfaces meniscus implants were removed fromtheir original sterile packaging and assigned randomly intothe two groups. The initial weight of the implants wasmeasured at the starting point without prior heating ordrying— as they were removed from their original packaging.The implants to be tested under static conditions were thensoaked in SPF (37 1C) and their weights were recorded again 6,24, 48 h; 1, 2, 4 weeks; 3 and 6 months after placing thespecimens in SPF. Fluid absorption was calculated gravime-trically by subtracting the weight of each implant at everytime point from its initial weight, recorded at the baselinemeasurement, and presented as the weight difference in

Femur

Tibia

Medial meniscus implantLateral

meniscus

Load cell

Femur replica

Tibia replica

Meniscus Implant

Closed cell filled with SPF (simulated physiological fluid)

Fig. 1 – The NUsurface Meniscus Implant (a), illustration of the knee joint and the location of the meniscus implant (b), and thein-vitro experimental set-up of the medial compartment replicating the femur and tibia surfaces, joint fluid, and axialloading (c).

j o u r n a l o f t h e m e c h a n i c a l b e h a v i o r o f b i o m e d i c a l m a t e r i a l s 2 9 ( 2 0 1 4 ) 4 2 – 5 544

either [mg] or [%]. Specimens which were tested in thedynamic simulation were placed in SPF (37 1C) for two weeksprior to simulation, and weighed. The implants were weighedagain after 100,000; 200,000; 300,000; 500,000; and 700,000; and1, 1.5 and 2 million load cycles. Fluid absorption was calcu-lated as was described for the static test.

2.2. Strain rate measurements

The baseline strain-rate response of the implant in its initialstate without prior soaking was measured while fully sub-merged in SPF (37 1C). The specimens were placed in the uni-axial tension-compression testing machine (Tinius Olsen,H10KT), between rigid UHMWPE tibia and femur replicas(Fig. 1c). The replicas were manufactured by machine-millingtwo highly wear-resistant UHMWPE blocks (Polystones MR,Rochling, Germany) to the geometry of the articulating sur-faces of a size-matched knee joint which was reconstructedfrom MRI scans (Elsner et al., 2010a). A 10 N pre-load wasapplied to make sure that the implant was seated properly.The distance between the polyethylene loading platens (l0)was recorded in this state for future strain calculations.

A loading program based on the application of a max-imum load of 1500 N (corresponding to 60% of the peak value

specified in ISO-1423) was applied using the following rangeof strain rates: 0.5, 1, 5, 10, 50, and 100 mm/min which arewithin the range reported in similar studies and in the samerange of physiological loading rates (Chia and Hull, 2008;Li et al., 2003). The maximal load was applied on eachspecimen, each time using a different strain rate of the rangementioned before. Force–displacement response under eachof the loading rates was recorded. The measurements wererepeated every time fluid absorption was recorded: after 6, 24,48 h; 1, 2, 4 weeks; 3 and 6 months soaking in SPF, andafter 100,000; 200,000; 300,000; 500,000; 700,000; and 1, 1.5 and2 million load cycles of the dynamic test.

Due to the compliant nature of the implant, there was also aneed to characterize the evolution of the contact area — A(t),between the inferior surface of the implant and the tibialreplica as a function of the applied force. An implant wasplaced on the uni-axial compression testing machine betweenthe tibia and femur replicas with a piezoelectric pressuresensor (Tekscan Inc. Boston, MA) placed between the implantand the tibial surface as described previously by Linder-Ganzet al. (2010). The implant was subjected to a 1500 N force, at astrain-rate of 10 mm/min, while continuously recording thecontact area. The resultant contact areas A(t) versus forcemeasurements F(t) were then used to establish the relationship

j o u r n a l o f t h e m e c h a n i c a l b e h a v i o r o f b i o m e d i c a l m a t e r i a l s 2 9 ( 2 0 1 4 ) 4 2 – 5 5 45

between A(t), F(t), and momentary stresses. The relationshipsbetween force-deformation and stress–strain were calculatedusing two methods:

Method A: The stress was calculated by dividing the compres-sive force by the constant total (inferior) surface area of theimplant — A (1172mm2). This calculation method yields alower bound for the true stress, which would ideally occur ifall of the implant comes into contact with the tibial surface.Method B: The nominal stress was calculated by dividing thecompressive force by the momentary contact area, A(t). Thiscalculation method yields an upper bound for the true stresssince the pressure sensor underestimates the real contactarea due to a minimal threshold for detection of contact.

2.3. Creep measurements

The specimens were placed on the uni-axial compression testingmachine, between the tibial and femur replicas and a 10 N pre-load was applied to make sure the implant was seated inposition. A 10-step loading program based on rapid applicationsof loading phase was applied (150 N, and increasing by 150 Nincrements at each following step). A crosshead speed of100mm/min was set up to 90% of the desirable force (step value)followed by a 10mm/min compression rate which was appliedfor the 10% remaining force. This testing procedure was per-formed to avoid an overshoot of more than 10% of the forcevalue. Upon the constant loading phase, the compressive dis-placement was monitored as a function of time over 90 s. Theload was then released and kept close to zero (10 N) for another90 s to monitor the recovery of the deformation after the loading.After completing each of the tests, the implants were dried andmeasured in order to evaluate any permanent geometricalchanges of the width, length or thickness of the implant, causedeither due to soaking, or due to repeated loading.

2.4. Relaxation measurements

The specimens were placed as in the creep test on the testingmachine, and a 10 N pre-load was applied. A 10-step loadingprogram in which compressive deflection values were increasedby increments representing 10% themaximal deflection obtainedunder 1500 N was applied. The test specimens were compressedto the required deflection levels using a crosshead speed of1000mm/min. After the appropriate deflection step was reached,it was held constant and the decay in force was recorded as afunction of time, for a period of 90 s. The strain was thenreleased, kept at zero position for another 90 s, and the nextstep was applied and so forth.

2.5. Hysteresis measurements

Hysteresis was measured by subjecting the implant to 10consecutive loading–unloading cycles employing loadingrates of 1, 10, and 100 mm/min, between 10 and 1500 N.Force versus displacement measurements were recorded.This sequence was only repeated for the specimens in thedynamic simulation, at each of the aforementioned testingpoints specified in Section 2.1.

2.6. Analysis and statistics

Results were compared using a one-way-ANOVA followed by apost-hoc Tukey-Kramer test in cases where only one dependentvariable exists (e.g., the relaxation time constant versus priorsoaking time or loading cycles). In cases of dependence of twovariables (e.g., relative relaxation versus step, and versus soakingtime or loading cycles) a two way-ANOVA was conducted toevaluate the significance of interaction, followed by a post-hocTukey-Kramer test to evaluate significant effects in betweengroups.

3. Results

3.1. Fluid absorption

Implants which were placed in SPF were found to gain 62 mgper implant or 0.75% of their initial weight over six months.The most significant weight gain was reached at the firstweek with 0.65% weight gain. Thereafter, fluid absorptiononly slightly increased in the following weeks (Fig. 2a). For thesamples that were soaked in SPF for two weeks and thentested dynamically, the initial absorption was approximatelythe same (0.7%) but fluid absorption was found to increasefurther due to the loading, until stabilizing at approximately1.05% after 1.5 million cycles (Fig. 2b).

3.2. Strain rate

Force–displacement curves attained during the loading condi-tions mentioned above were normalized to stress and strainvalues once by using the total available surface area (‘MethodA’), and a second time, by using the developing contact areawhich was measured while subjecting the implant to 1500 Ncompression at 10mm/min rate (‘Method B’).

Mild changes in the stress–strain curves occurred as aresult of the increase in strain-rate. The maximum strainmeasured for the virgin implant under 1500 N load was8075% at 1 mm/min, 7672% at 10 mm/min, and 7474% at100 mm/min. After six months soaking in SPF these valueswere reduced to 7572%, 7172% and 6970%, respectively(pr0.05). The effect of dynamic loading was even moreapparent, with a maximum strain of 5974%, 5573% and5474% (pr0.05).

The evolution of the contact area against the applied forcewas found to be non-linear (Fig. 3a and specific examplesof the contact area at 500, 1000 and 1500 N in Fig. 3b) andtherefore the curve was fitted to a four-degree polynomialfunction (R2¼0.999):

CA¼�6:88� 10�10 � F4 þ 2:38� 10�6 � F2�3� 10�2 � F2

þ2:2� Fþ 100:52 ð1Þwhere F is the applied force and CA is the correspondingcontact area measured by the pressure sensor.

Overall, the elastic modulus increased mildly followinginitial soaking, but then remained near constant over time.The resultant stress–strain curves for three strain rates (slow,medium, and fast): 1, 10, and 100 mm/min are presented inFig. 3c and d for ‘Method A’ and ‘Method B’, respectively. Both

0

0.2

0.4

0.6

0.8

1

1.2

1.4

0 0.2 0.4 0.6 0.8 1

Stre

ss [M

Pa]

Strain

1 mm/min

10 mm/min

100 mm/min

0

0.2

0.4

0.6

0.8

1

1.2

1.4

0 0.2 0.4 0.6 0.8

1

Stre

ss [M

Pa]

Strain

1 mm/min

10 mm/min

100 mm/min

0

200

400

600

800

1000

1200

1400

0 500 1000 1500 2000

Con

tact

are

a [m

m2 ]

Force [N]

1500 N500 N 1000 N

12.211.1108.97.86.75.64.43.32.21.10.1MPa

Phase I

Phase II

Phase III

13.4

Fig. 3 – (a) Contact area measurements recorded by pressure sensors placed between the implant and a replica of the tibialplateau when a 1500 N load was applied axially on the implant at a strain rate of 10 mm/min. (b) Snapshots of the pressuremaps attained at 500, 1000 and 1500 N demonstrate the evolution in the contact area. (c) Stress–strain curves calculatedby using the total inferior surface area of the implant, and (d) true (measured) contact area of the implant at 1, 10 and100 mm/min strain-rates.

0.0

0.2

0.4

0.6

0.8

1.0

1.2

1.4

0 30 60 90 120 150 180

Diff

eren

ce f

rom

initi

al w

eigh

t [%

]

Time [days]

0.0

0.2

0.4

0.6

0.8

1.0

1.2

1.4

0 0.5 1 1.5 2

Diff

eren

ce f

rom

initi

al w

eigh

t [%

]

Time [x106cycles]

Fig. 2 – Fluid absorption (%) over time, (a) static soaking in simulated physiological fluid (SPF), and (b), following simulatedphysiological loading over 2 million cycles.

j o u r n a l o f t h e m e c h a n i c a l b e h a v i o r o f b i o m e d i c a l m a t e r i a l s 2 9 ( 2 0 1 4 ) 4 2 – 5 546

types of curves indicate stiffening of the implant underhigher loads, but the effect is more pronounced whenstresses are calculated according to Method A. Under higher

loads, the implant material conforms to the tibial surface, toutilize more of the theoretical surface area, and the curvesconverge to display similar stress–strain relationship.

j o u r n a l o f t h e m e c h a n i c a l b e h a v i o r o f b i o m e d i c a l m a t e r i a l s 2 9 ( 2 0 1 4 ) 4 2 – 5 5 47

The curves shown in Fig. 3c share a similar pattern that iscomposed of three distinct phases:

Fig2 manthemm

Phase I: A linear phase in which the implant has a relativelylow elastic modulus. This phase is seen between verticaldeflection values of 0–0.7mm (equivalent to approximately0–20% strain).Phase II: A non-linear phase of gradual stiffening, found tooccur between 0.7–1.7 mm deflections (equivalent toapproximately 20–50% strain).Phase III: A second linear phase in which the implantdisplays high resistance to strain compared to that recordedin phase I (50–80% strain).

The elastic moduli of the implant were not found to beaffected significantly by static soaking, nor by repeated load-ing: The 2% modulus (Phase I) was found to be inthe ranges of 0.13–0.20 MPa and 0.12–0.14 MPa for staticand dynamic soaking conditions, respectively, according toMethod A (Fig. 4a), and 0.92–1.31 MPa and 0.89–1.03 MPa,respectively, according to Method B (Fig. 4b). The modulus at

0

0.2

0.4

0.6

0.8

1

1.2

1.4

0 30 60 90 120 150 180

mod

ulus

(0.0

2 st

rain

) [M

Pa]

Soaking time [days]

0

1

2

3

4

5

6

0 30 60 90 120 150 180

mod

ulus

(0.6

6 st

rain

) [M

Pa]

Soaking time [days]

. 4 – Evolution of the elastic modulus following static and dyillion load cycles, respectively. The moduli were measured a

d 0.6 (c, d) strain for the slowest and fastest strain rates (1 mtheoretical contact area (Method A) a, (▲ — 1 mm/min, ● — 1/min, ◯ — 100 mm/min) are presented per case.

60% strain (Phase III) was found to be in the ranges of 3.79–4.54 MPa and 4.61–5.21 MPa for static and dynamic conditions,respectively, according to Method A (Fig. 4c) and 2.02–2.53, and2.61–3.08 MPa, respectively, according to Method B (Fig. 4d).

3.3. Creep

The time dependant creep response was measured by apply-ing incremental step loads on to the implant. The instanta-neous and asymptotic strains per load are shown for thevirgin implants, and for the implants which underwent 2million fatigue cycles prior to testing (Table 1). The creepresponse of the virgin implant and of an implant soaked inSPF for 6 months is shown in Fig. 5a. The creep response of apre-soaked, unloaded implant and that of an implant whichunderwent 2 million fatigue cycles is shown in Fig. 5b.

As expected according to the theory of viscoelastic materials(Ward and Sweeney, 2004), the strain response is composed of3 distinct phases:

i.

0

0

0

0

1

1

mod

ulus

(0.0

2 st

rain

) [M

Pa]

mod

ulus

(0.6

str

ain)

[MPa

]

namt thm/m00

ε1 — an immediate strain response depicted by an instantincrease (during loading) or decrease (during unloading)

0

.2

.4

.6

.8

1

.2

.4

0 1000 2000Loading time [ 1000 cycles]

0

1

2

3

4

5

6

0 1000 2000Loading time [ 1000 cycles]

ic soaking conditions for the durations of 6 months ande linear portions of the stress–strain curves at 0.02 (a, b)in and 100 mm/min). Results for calculations based onmm/min) and effective contact area (Method B), (Δ — 1

Table 1 – Applied loads and resultant instant and symptomatic strain values, and relative change in the strain (creep) asmeasured for virgin implants and implants which previously underwent 2 million fatigue load cycles in SPF.

Step Appliedforce (N)

Virgin implant Implant after 2 million cycles

Instant strain Asymptotic strain Relativecreep (%)

Instant strain Asymptotic strain Relativecreep (%)

1 150 0.2470.01 0.2870.01 17.5373.09 0.2270.02 0.2570.12 14.1973.272 300 0.3470.01 0.3970.05 16.2979.94 0.2970.03 0.3370.04 13.2571.873 450 0.4070.01 0.4570.01 11.0771.33 0.3670.04 0.4170.05 12.3771.014 600 0.4670.01 0.5170.01 10.5771.39 0.4270.05 0.4770.05 11.7270.575 750 0.5070.01 0.5670.01 10.5971.80 0.4670.05 0.5170.05 10.6470.536 900 0.5570.00 0.6070.01 9.9071.75 0.4970.05 0.5570.05 11.2971.217 1050 0.5870.01 0.6470.01 9.8471.98 0.5470.05 0.5870.05 8.8871.198 1200 0.6270.00 0.6870.01 9.9271.19 0.5670.05 0.6170.05 9.6070.529 1350 0.6570.01 0.7170.01 9.5472.00 0.5970.05 0.6470.05 8.1070.29

10 1500 0.6870.01 0.7570.01 9.6871.24 0.6170.05 0.6770.06 10.0571.49

Data shows means7standard deviations.

j o u r n a l o f t h e m e c h a n i c a l b e h a v i o r o f b i o m e d i c a l m a t e r i a l s 2 9 ( 2 0 1 4 ) 4 2 – 5 548

in the position of the compression jig in response to theload step.

ii.

ε2 — further gradual increase or decrease in the position ofthe compression jig, while the load is maintained constant.This response fades over time, and the time constant (τ) forthis decay is determined by the time needed for it to decayto approximately 37% of the difference between its initialand asymptotic value. The τ values were calculated forboth loading (τ1) and unloading (τ2) phases.

iii.

ε3 — residual strain which can be detected by a gradualincrease in the baseline position of the compression jigover time. The lower most point of the unloading responsetends to increase from one loading cycle to the next.

Average values of ε1�ε3 at 300, 600 and 1200 N can befound in Table 2. While ε3 increases linearly, neither ε1 nor ε2comply with the rule of linearity, and thus overall, thematerial behaves as a non-linear viscoelastic material.

Throughout the steps, the relative creep was found to beapproximately 10% of the initial strain experienced by theimplant when the loading step was applied. The ANOVA testshowed dependence of the relative creep on the loading step,and the post-hoc Tukey Kramer test confirmed that for thefirst two loading steps (150 and 300 N) this ratio was higher(�17%) than the following steps (pr0.05). Nevertheless, nostatistical differences were found between loading steps 3–10.The ANOVA also showed no statistical difference betweenthe creep response of the virgin implants and implants whichunderwent 2 million fatigue cycles. Similarly, the creep decaytime constants (τ's) were found to be near constant betweenloading steps. The averaged decay time constant τ was 18.770.3 s for a virgin implant, but increased gradually by 40% onaverage to 2570.4 s as a result of exposure of the implant toSPF (pr0.05). In the dynamic test, the decay time constantafter 2 million load cycles was also significantly higher thanthe virgin implant by 40% (2571.5 s, pr0.05), but was notdifferent from the τ measured after 6 month's static soaking.The evolution in the values of τ as a function of staticsoaking and dynamic loading is presented in Fig. 5c and drespectively.

Long-term creep was evaluated by measurements of theimplants' gross width, length and thickness. The initial mea-surements of a virgin implant were 37.3170.07, 49.9570.22and 13.1370.03mm, respectively. Statistical tests of signifi-cance between implants in each test group showed thatimplants subjected to 6 months' soaking had not undergonesignificant changes in width (37.0370.13mm, �0.8%) or length(49.9070.10mm, �0.1%) compared to the virgin implant, butdid undergo a slight increase in thickness (13.2170.07mm,þ0.6%), which was statistically significant (pr0.05). On theother hand, implants from the group subject to fatigue loadingover 2 million cycles had increased slightly in both width(37.6670.41mm, þ0.9%) and length (50.5270.36mm, þ1.1%),but decreased in thickness (13.0070.11mm, �1%), all of whichwere found to be statistically significant.

3.4. Relaxation

Relaxation was measured by applying deflection steps whichcorrespond to the force steps applied during the creep mea-surements, for a duration of 90 s, and releasing them for thesame time. The instantaneous and asymptotic loads mea-sured under these steps are shown in Table 3. A representativestress relaxation response of an implant placed in SPF for 6months, and its control (virgin implant) is shown in Fig. 6a.Similarly, the response of an implant following 2 millionfatigue cycles and its control (pre-soaked, unloaded implant)is shown in Fig. 6b. The response typically consists of twophases; (i) an instant increase in force, followed by (ii) gradualdecay in the force value until reaching an asymptotic value.Larger variability was seen in the relative relaxation in loadbetween steps compared to creep. The relative relaxation inload was found to be 21.875.5% (Table 3), and only differentsignificantly in the last loading step (pr0.05). After 2 millioncycles, the relative relaxation remained similar 22.778.0% andwas not statistically different compared to the virgin state.

As indicated for the creep measurements, the relaxationresponse was found to fade over time, and the time constant (τ)for this decay is determined by the time needed for it to decayto approximately 37% of the difference between its initial(instant) response and the asymptotic value. The evolution in

0

0.1

0.2

0.3

0.4

0.5

0.6

0.7

0.8

0 500 1000 1500 2000

Stra

in

Time [Seconds]

0

0.1

0.2

0.3

0.4

0.5

0.6

0.7

0.8

0 500 1000 1500 2000

Stra

in

Time [Seconds]

0

5

10

15

20

25

30

0 500 1000 1500 2000

Cre

ep T

ime

Coe

ffici

ent [

sec

]

Soaking time [ x 1000 cycles ]

0

5

10

15

20

25

0 30 60 90 120 150 180

Cre

ep T

ime

Coe

ffici

ent

[sec

]

Time [Days]

Fig. 5 – (a) Representative creep response of specimens from the static soaking test from a virgin implant (●) and an implantsoaked in simulated physiological fluid (SPF) for 6 months ( ). (b) Representative creep response of specimens from thedynamic soaking test from a implant which was pre-soaked for two weeks (●) and an implant which underwent 2 million loadcycles in SPF ( ) (c) Evolution of the average (7standard deviation) creep time coefficient τ as measured during the staticsoaking test (d) Evolution of the average (7standard deviation) creep time coefficient τ as measured during the dynamic test.

Table 2 – Implant strain values calculated from creep tests of a wet control under specific loads.

Immediate elastic deformation (e1) Delayed elastic deformation (e2) Newtonian flow (e3) Total strain

300 [N] 0.3170.02 0.0570.01 0.0170.00 0.37870.135600 [N] 0.4370.02 0.0670.01 0.0370.01 0.52470.1841200 [N] 0.5870.02 0.0870.01 0.0570.01 0.71170.244

j o u r n a l o f t h e m e c h a n i c a l b e h a v i o r o f b i o m e d i c a l m a t e r i a l s 2 9 ( 2 0 1 4 ) 4 2 – 5 5 49

τ values measured during the static and dynamic tests areshown in Fig. 6c and d respectively. The baseline τ calculated forrelaxation for a virgin implant was 21.171.9 s. Soaking in SPFfor 6 months had little effect on the relaxation properties of theimplants with no statistical difference from a virgin implant.Looking at the averaged relaxation time coefficients, it can beseen that the mean decay time constant varied little over time,and after 6 months soaking in SPF is 21.471.6 s. The relaxationtime coefficient calculated from this response after 2 millioncycles decreased slightly to 19.471.7 s compared to the baselinemeasurement prior to simulation (pr0.05).

3.5. Hysteresis

The finding of hysteresis loops supports the previousfindings that the implant behaves as a viscoelastic

material, as it dissipates energy and has a hysteresis loopin cyclic loading and unloading. Representative hysteresisloops attained for a virgin implant at the start of the testand following 2 million fatigue cycles are found in Fig. 7aand b, respectively.

Pre-conditioning was found to occur during the first andsecond loading–unloading cycles, but then loading andunloading pathways were found to repeat for the subsequentloading cycles (3–10). The hysteresis energy, as calculated forcycles 3–10, was found to be affected by the strain rate whenit was increased from 1 to 10 mm/min, and 1 to 100 mm/min(pr0.05) but not between 10 and 100 mm/min (p¼0.45).Despite a trend of decrease in the hysteresis energy up to300,000 load cycles (Fig. 7c) the implant showed no depen-dence between the hysteresis energy and fatigue loading upto 2 million cycles (p¼0.053).

Table 3 – Applied loads and resultant instant and symptomatic strain values, and relative creep as measured for virginimplants and implants which previously underwent 2 million fatigue load cycles in SPF.

Step Deformation(mm)

Virgin Implant Implant after 2 million cycles

Maximumload (MPa)

Asymptoticload (MPa)

Loadrelaxation(%)

Maximumload (MPa)

Asymptoticload (MPa)

Loadrelaxation(%)

1 0.25 0.1670.00 0.1370.00 22.4371.61 0.1770.01 0.1470.01 18.6371.662 0.5 0.2470.01 0.1970.00 27.3472.71 0.2670.02 0.2170.02 18.9771.623 0.75 0.3270.01 0.2770.01 20.5871.83 0.3670.02 0.3070.02 15.8971.924 1 0.4270.01 0.3670.01 18.2070.96 0.4870.04 0.4170.03 15.4071.305 1.25 0.5470.01 0.4670.01 19.0470.82 0.6170.05 0.5170.04 16.3571.326 1.5 0.6670.01 0.5670.01 18.2370.43 0.7470.06 0.6170.05 17.8071.707 1.75 0.8170.02 0.6770.01 21.0270.77 0.8870.09 0.7370.07 17.3570.788 2 0.9770.02 0.8170.01 19.2671.18 1.0170.08 0.8570.09 15.3672.069 2.25 1.1070.01 0.9670.02 14.2571.60 1.1570.11 0.9770.09 15.5772.43

10 2.5 1.4170.08 1.0870.01 31.4075.01 1.5770.35 1.0870.08 30.96711.04

Data shows means7standard deviations.

0

0.2

0.4

0.6

0.8

1

1.2

1.4

1.6

0 500 1000 1500 2000

Stre

ss [

MPa

]

Time [ seconds ]

0

0.2

0.4

0.6

0.8

1

1.2

1.4

1.6

0 500 1000 1500 2000

Stre

ss [

MPa

]

Time [ seconds ]

0

5

10

15

20

25

30

0 500 1000 1500 2000

Rel

axat

ion

Tim

e co

effic

ient

[ se

c ]

Soaking time [ x 1000 cycles ]

0

5

10

15

20

25

30

0 30 60 90 120 150 180

Rel

axat

ion

tim

e co

effic

ient

[ se

c ]

Soaking time [days]

Fig. 6 – (a) Stress relaxation responses of specimens from the static soaking test from a virgin implant (●) and an implant soaked insimulated physiological fluid (SPF) for 6 months ( ). (b) Representative stress relaxation responses of specimens from the dynamicsoaking test from an implant which was pre-soaked for two weeks (●) and an implant which underwent 2 million load cycles in SPF( ). (c) Evolution of the average (7 standard deviation) relaxation time coefficient τ as measured during the static soaking test (d)Evolution of the average (7standard deviation) relaxation time coefficient τ as measured during the dynamic test.

j o u r n a l o f t h e m e c h a n i c a l b e h a v i o r o f b i o m e d i c a l m a t e r i a l s 2 9 ( 2 0 1 4 ) 4 2 – 5 550

0

0.2

0.4

0.6

0.8

1

1.2

1.4

1.6

0 0.2 0.4 0.6 0.8

Stre

ss [M

Pa]

Strain

0

0.2

0.4

0.6

0.8

1

1.2

1.4

1.6

0 0.2 0.4 0.6 0.8

Stre

ss [M

Pa]

Strain

0

0.1

0.2

0.3

0.4

0.5

0.6

0.7

0.8

0.9

1

0 500 1000 1500 2000

Ener

gy [

MPa

]

1000 Fatigue cycles

1mm/min

10mm/min

100mm/min

Fig. 7 – Representative hysteresis loops as measured for avirgin implant (a) and after 2 million fatigue cycles (b) byapplying 10 consecutive 100 mm/min strain-rate loading(1500 N) —unloading cycles. (c) Trends of the hysteresisenergy over time, as a function of the applied strainrate:� �— 100 mm/min, ▲ — 10 mm/min, ● — 1 mm/min.

j o u r n a l o f t h e m e c h a n i c a l b e h a v i o r o f b i o m e d i c a l m a t e r i a l s 2 9 ( 2 0 1 4 ) 4 2 – 5 5 51

4. Discussion

In this study, the viscoelastic properties of a novel meniscusimplant were determined. Specifically, the study includedcompression tests at different strain rates, step stress/strainresponse tests (creep and relaxation), and hysteresis tests.The experiments were conducted on ‘virgin’ implants,implants that underwent short and long-term soaking, aswell as devices that underwent 2 million fatigue loadingcycles. In general, the meniscus implant was found to main-tain its overall structure throughout the testing and demon-strated a typical viscoelastic behavior, similar to the naturalmeniscus (Chia and Hull, 2008). This finding was shownclearly by the typical stress–strain curves, tangent elasticmodulus results, strain rate-dependent behavior, creepresponse, relaxation pattern, and hysteresis. Other than arecent three-dimensional finite element knee model whichwas used to compare the creep behaviors of normal andmeniscectomized knees (Kazemi et al., 2011), the viscoelasticbehavior of the meniscus has been investigated using smallsections of the material rather than in the knee joint setting(Uezaki et al., 1979; Chia and Hull, 2008; Bursac et al., 2009).The method used in this study is unique in that it utilizes acomplete meniscus and a knee model to characterize themechanical properties of the meniscus under compression.

The overall structure or appearance of the device showedno changes following static soaking in SPF for the durationof 6 months or following 2 million dynamic loading cycles.Static soaking in SPF resulted in a 0.75% weight gaindue to fluid absorption over 6 months; whereas cyclic fatigueincreased the absorption further to 1.05% (Fig. 2). Moistureabsorption behavior depends on the material's initial fluidcontent, temperature, applied load, type of media, time andmaterial system (Oliveira and Creus, 2004). However theresults are similar to the hygroscopic capacity of PCU (Bio-nates 80A) which was reported by Khan et al., (2005) andGeary et al. (2008) (0.8–1.3%). The finding that the implant isable to further absorb fluid under dynamic loading conditionshas also been described in another long-term wear study ofan acetabular hip component composed of PCU (Elsner et al.,2010b, 2011). These results imply that the PCU material haslimited availability of sites for hydrogen bonding of watermolecules compared to other materials which have beenused as meniscus implants (Kobayashi et al., 2005; vanTienen et al., 2002). Since the overall fluid contents of theimplant are significantly lower than that of the naturalmeniscus (�72%, Herwig et al., 1984; Allen et al., 1995), fluidpressurization and flow are not expected to have a dominanteffect on viscoelasticity, and consequently, the viscoelasticbehavior observed in this study is determined by the intrinsicproperties of the implant materials. The NUsurfaces implantis composed mainly (over 95% weight) of PCU, but it alsoincorporates UHMWPE reinforcement in the form of intern-ally embedded fibers. As a composite structure, it is expectedto demonstrate a combined effect of these elements based ontheir relative amounts and individual properties. The typicalstress–strain curves attained for the implant during compres-sion, exhibit this type of behavior, resulting in three phases(Fig. 3c,d):

Phase I: Under low compressive loads, the implant is fairlyflexible, and it is able to deform easily. Such behavior is

j o u r n a l o f t h e m e c h a n i c a l b e h a v i o r o f b i o m e d i c a l m a t e r i a l s 2 9 ( 2 0 1 4 ) 4 2 – 5 552

governed by the mechanical properties of the main com-ponent, PCU.Phase II: As the compressive load applied on the implantis increased, the implant gradually stiffens, and resistsdeformation. This change can be attributed to the pre-sence of the reinforcing UHMWPE fibers which are slack intheir native state, but gradually become aligned andtensed as the implant expands radially.Phase III: As the implant is compressed and undergoes radialexpansion, some of the reinforcement fibers straightenbefore others, and thus the stiffness of the sample increasesgradually and becomes constant when all the fibers becometensed.

A similar mechanical response has been reported for thenatural meniscus tissue by Tissakht and Ahmed (1995) andChia and Hull (2008), with the difference being that for thenatural meniscus tissue, the hydrated extracellular matrixacts as the elastic material, whereas the collagen fibers act asthe reinforcing fibers.

The characterization of the gross mechanical response ofthe meniscus (as opposed to a standard plug/strip section) ischallenging due to changes in the implant's contact area withthe adjacent surfaces during loading (Fig. 3). This is furthercomplicated by its non-uniform thickness. Consequently, twocalculation approaches were used to calculate its elasticmodulus under compression: (A) applied force/total surfacearea (constant value), and (B) applied force/momentary con-tact area (measured), provide a lower value to the estimationof stress. Thickness was selected as the minimum in bothcases. The resultant moduli calculated in both methods werefound to be in the range of 100 kPa and 1.3 MPa in the 2%region, which agrees with the compressive modulus reportedby Chia and Hull (2008) for sections of the natural meniscus.

Another point of similarity to the mechanical behavior ofthe native meniscus is that the implant exhibits higherresistance to deformation under higher compression rates.Strain rate dependence occurs because slower strains allow forgreater relaxation to take place within the tested material.This relaxation allows time for the polymer chains to reconfi-gure and slide upon each other (Fung, 1993; Ward andSweeney, 2004). For higher strain rates, polymer chains gettangled with each other and therefore maximal strain at agiven stress is smaller. Higher strain rate demonstrates con-tinuous increase in stiffness and reduction in the maximumstrain under the maximal load (1500 N) with increasing rate(Fig. 3). When evaluating the strain rate response of theimplant as a function of exposure to simulated body fluid,changes to the maximum strain and moduli were found to bevery mild (Fig. 3). Generally, there is a trend towards reductionin the maximum strain or increased stiffness after soaking,compared to the baseline measurement. This is a surprisingfinding since previous studies have reported that unreinforcedPCU tends to soften and lose up to 8% of its initial elasticmodulus under hygroscopic aging conditions due to plastici-zation (Khan et al., 2002; Geary et al., 2008). Nevertheless, thecurrent study of compressive loading of a reinforced meniscusimplant differs considerably from, and appears to be more

relevant than, these tensile tests, which were based onmeasure-ments of thin homogenous PCU sheets.

When subjected to step stress (creep) or strain (relaxation),the meniscus implant displayed viscoelastic behavior, whichqualitatively resembles that of the natural meniscus. Under aconstant stress/force, the meniscus displays an initial phase ofimmediate elastic response, followed by a transient deformationwhich stabilizes within 100 s from applying the load (Joshi et al.,1995; Chia and Hull, 2008; Kazemi et al., 2011). In the naturalmeniscus, the immediate response is governed by the hydro-static pressure developed in the fluid portion of the tissue. This isfollowed by continuous deformation of the tissue at a slower rateas the fluid is expelled from the matrix, with the solid matrixresisting more of the load (Athanasiou, 2009). A similar creepresponse, consisting of an immediate deformation followed by adelayed response, was demonstrated by the implant. However,while the viscoelastic behavior of the natural meniscus isattributed to its interplay between solid and fluid phases, theviscoelastic behavior of the implant is attributed to the compo-site solid structure mainly due to the PCU component, which isrestrained by the limited radial extension of the reinforcingUHMWPE fibers (Elsner et al., 2010a). Viscoelasticity is evidentin the virgin state, when the implant has not yet absorbed anyfluid, and does not change considerably even when fluidabsorption has peaked (Fig. 5). In polymeric materials, the creepresponse is attributed to the molecular motion in the solidinduced by the stress. Parts of the polymer's chains changeposition and rearrange in order to comply with the stressapplied. This creates a back pressure in the material and onlywhen the back pressure is the same magnitude as the appliedforce, the material no longer creeps. After the load is removed,the chains return to their equilibrium conformation due to anentropy-driven tendency to restore the initial state (McCrumet al., 2003). It is noteworthy that the viscoelastic response of theimplant is not linear because PCU is hyper-elastic and does notbehave linearly when subjected to large deformations as itexperienced in this study (Elsner et al., 2010a).

The creep behavior of the implant is of relevance to itsshort-term and long-term function under physiological condi-tions. The typical time constant for the delayed creep responsewas found to be between 20 and 25 s; suggesting that theimplant is not likely to experience a full creep response whenthe maximum load is applied on it during gait conditions.Similar values were found in both soaked and loaded implants,e.g., static soaking in SPF was found to increase the timeconstant from 20 to 25 s, but for samples which were presoakedand underwent dynamic fatigue simulation, the decay timeconstant remained approximately 25 s.

In the long term, as seen after 2 million load cycles, theimplant's width and length increased slightly (0.9% and 1.1%,respectively) and thickness reduced (�1%) compared to itsinitial state (pr0.05). Permanent creep due to continuousmechanical loading is attributed to any irreversible deforma-tion of the implant's constituent components. The UHMWPEfibers are not predicted to creep considerably due to theirhigh modulus (140 GPa) and their low elongation to break(2.9–4.1%) (DSM Puritys, product datasheet, 2013), thereforethe long-term residual creep of the implant most likelyreflects permanent deformation of the PCU matrix. Highlylocalized deformation zones often lead to irreversible sliding

j o u r n a l o f t h e m e c h a n i c a l b e h a v i o r o f b i o m e d i c a l m a t e r i a l s 2 9 ( 2 0 1 4 ) 4 2 – 5 5 53

of polymer chains. A similar phenomenon of plastic deforma-tion of PCU was previously reported by Cipriani et al. (2013)for PCU-based spine spacers. Moderate creep of the PCU bulkunder gait conditions could be considered as an advantage inan implant which is expected to articulate between existingbiological surfaces. The implant can adjust itself to variationsin the joint morphology of a specific patient, increase con-formity, and improve pressure distribution.

The stress-relaxation tests generally reconfirm the resultsof the creep characterization. The immediate stress responsewas followed by a continuous exponential decay of the appliedstress, and stabilized within 100 s. The decay time constant τ

of approximately 21 s is within the above mentioned rangereported for creep, and slightly longer than the viscoelasticbehavior of natural meniscus and cartilage (10–13 s; Suh andDiSilvestro, 1999; June et al., 2009). Although static soaking alonedid not affect the relaxation time coefficient over 6 months (21 s),repeated loading resulted in a significant decrease of �10% in itsvalue, to approximately 19 s (pr0.05). However, the differencebetween these values is unlikely to have any clinical relevance.

Degradation of the implant's structure connected withfatigue, involves some changes in parameters of the hyster-esis loop that can usually forecast material fatigue failure.When subjected to cyclic loading, the implant demonstrateda hysteresis response which follows the typical hysteresispattern of soft tissues (Park et al., 2004, Schatzmann et al.,1998). The last consists of large hysteresis value at the firstcycle, followed by the dissipation of the kinetic energy over aloading cycle, until it reaches a steady state. However, whilepre-conditioning of the implant it was found to occur duringthe first and second loading–unloading cycles only, thenumber of cycles to achieve a reproducible hysteresis loopof a cartilaginous tissue is typically one order of magnitudehigher (Kelm, 2002). This time-dependent behavior of theimplant, which reaches a steady-state almost immediatelywhen cyclic loading is applied, is important during dailyactivities, such as walking, running, and cycling. The hyster-esis energy was found to decrease when the loading rateincreased. However, the location of the loop was stablethroughout the dynamic test, thus providing further assur-ance that the combined effects of soaking and loading on thestability of the implant are negligible.

Several limitations associated with the test model andassumptions of the present study deserve further discussion.Only axial loading was applied to the test specimens in thedynamic test, while ignoring the effect of shear loading as aresult of articulation. Since the axial load is the dominant loadin the knee, and it was found to have little effect on the long-term properties of the device, it is not expected that lower shearforces would result in significant changes to the results. How-ever, further testing using full mixed-mode loading conditionsmay provide further simulation of the clinical use of theimplant. Another potential limitation of the test set-up is theuse of polyethylene replicas for the cartilaginous surfaces.Although these were manufactured to match the geometry ofthe natural tibia and femur, they are significantly more rigid(Shore hardness 65D) and show lesser fluid absorption (o0.01%)and thus allow for higher local stresses to develop at theinterface. Based on this, this test set-up likely represents aworst-case condition to clinical use.

While the use of whole-joint testing, compared to smallsample testing, is advantageous in representing the real jointloading conditions, translating the raw data into standardstress and strain values was challenging due to the non-uniform geometry of the implant. Different approaches wereused in these analyses and the results did seem to be sensitiveto the approach chosen in these calculations. Strains andstresses were found to remain in the same ranges.

5. Conclusions

The meniscus implant was found to behave as a non-linearviscoelastic material, with mechanical properties similar tothose of the natural meniscus. The implant was foundto maintain its geometry when subjected to soaking in SPF.The combination of soaking, together with fatigue loadingresulted in mild geometrical changes as a result of creep. Thesechanges, in the order of magnitude of a fraction of millimeter,can be considered as the long-term adaptation of the implantunder load. All of the mechanical characterization tests showeda mild transition in the mechanical properties during the first300,000 load cycles, which then stabilized for the rest of theduration of 2 million cycles. These relatively small changes ingeometry reflect the adaptation measured after 2 million cyclesand appear to represent the effects of long-term use.

Acknowledgment

This study was funded by Active Implants Corporation(Memphis, TN, USA). M. Shemesh, R. Asher, E. Zylberberg,E. Linder-Ganz, and J.J. Elsner are employees of Active ImplantsCorp.

r e f e r e n c e s

Ahmed, A.M., Burke, D.L., 1983. In-vitro measurement of staticpressure distribution in synovial joints—Part I: tibial surface ofthe knee. ASME Journal of Biomechanical Engineering 105,216–225.

Allen, A.A., Caldwell, G.L., Fu, F.H., 1995. Anatomy andbiomechanics of the meniscus. Operative Techniques inOrthopaedics 5 (1), 2–9.

Allen, P.R., Denham, R.A., Swan, A.V., 1984. Late degenerativechanges after meniscectomy. factors affecting the knee afteroperation. Journal of Bone and Joint Surgery: British Volume66, 666–671.

Andrews, S., Shrive, N., Ronsky, J., 2011. The shocking truth aboutmeniscus. Journal of Biomechanics 44 (16), 2737–2740.

Athanasiou, K.A., Sanchez-Adams, J., 2009. Engineering the kneemeniscus, Edited by Athanasiou K.A., Synthesis Lectures onTissue Engineering, #1, Morgan & Claypool, San Rafael, CA.

Ateshian, G.A., 2009. The role of interstitial fluid pressurization inarticular cartilage lubrication. Journal of Biomechanics 42 (9),1163–1176.

Baker, P., Coggon, D., Reading, I., Barrett, D., McLaren, M.,Cooper, C., 2002. Sports injury, occupational physical activity,joint laxity, and meniscal damage. Journal of Rheumatology29 (3), 557–563.

Bedi, A., Kelly, N.H., Baad, M., Fox, A.J., Brophy, R.H., Warren, R.F.,Maher, S.A., 2010. Dynamic contact mechanics of the medial

j o u r n a l o f t h e m e c h a n i c a l b e h a v i o r o f b i o m e d i c a l m a t e r i a l s 2 9 ( 2 0 1 4 ) 4 2 – 5 554

meniscus as a function of radial tear, repair, and partialmeniscectomy. Journal of Bone and Joint Surgery: AmericanVolume 92 (6), 1398–1408.

Bullough, P.G., Munuera, L., Murphy, J., Weinstein, A.M., 1970. Thestrength of the menisci of the knee as it relates to their finestructure. Journal of Bone and Joint Surgery: British Volume52B, 564–570.

Bursac, P., Arnoczky, S., York, A., 2009. Dynamic compressivebehavior of human meniscus correlates with its extra-cellularmatrix composition. Biorheology 46 (3), 227–237.

Chia, H.N., Hull, M.L., 2008. Compressive moduli of the humanmedial meniscus in the axial and radial directions atequilibrium and at a physiological strain rate. Journal ofOrthopaedic Research 26 (7), 951–956.

Cipriani, E., Bracco, P., Kurtz, S.M., Costa, L., Zanetti, M., 2013.In-vivo degradation of poly(carbonate-urethane) based spineimplants. Polymer Degradation and Stability 98 (6), 1225–1235.

DSM/Dyneema Puritys website: ⟨http://www.dsm.com/en_US/medical/public/home/downloads/publications/Dyneema_Purity_UG_VG_fiber.pdf?fileaction=openFile⟩.

El Fray, M., Altstädt, V., 2003. Fatigue behaviour of multiblockthermoplastic elastomers. 2. Dynamic creep of poly(aliphatic/aromatic-ester) copolymers. Polymer 44, 4643–4650.

Elsner, J.J., Portnoy, S., Zur, G., Guilak, F., Shterling, A., Linder-Ganz, E., 2010a. Design of a free-floating polycarbonate-urethane meniscal implant using finite element modeling andexperimental validation. Journal of Biomechanics Eng. 132 (9),095001.

Elsner, J.J., Mezape, Y., Hakshur, K., Shemesh, M., Linder-Ganz, E.,Shterling, A., Eliaz, N., 2010b. Wear rate evaluation of a novelpolycarbonate-urethane cushion form bearing for artificial hipjoints. Acta Biomaterialia 6 (12), 4698–4707.

Elsner, J.J., Shemesh, M., Mezape, Y., Levenshtein, M., Hakshur, K.,Shterling, A., Linder-Ganz, E., Eliaz, N., 2011. Long-termevaluation of a compliant cushion form acetabularbearing for hip joint replacement: a 20 million cycleswear simulation. Journal of Orthopaedic Research 29 (12),1859–1866.

Englund, M., Guermazi, A., Gale, D., Hunter, D.J., Aliabadi, P.,Clancy, M., Felson, D.T., 2008. Incidental meniscal findings onknee MRI in middle-aged and elderly persons. New EnglandJournal of Medicine 359 (11), 1108–1115.

Fithian, D.C., Kelly, M.A., Mow, V.C., 1990. Material properties andstructure-function relationships in the menisci. ClinicalOrthopaedics and Related Research 252, 19–31.

Fung Y.C. Biomechanics. Mechanical properties of living tissues.2nd Ed. Springer, New York, USA, 1993.

Geary, C., Birkinshaw, C., Jones, E., 2008. Characterisation ofBionate polycarbonate polyurethanes for orthopaedicapplications. Journal of Materials Science: Materials inMedicine 19 (11), 3355–3363.

Hannink, G., van Tienen, T.G., Schouten, A.J., Buma, P., 2011.Changes in articular cartilage after meniscectomy andmeniscus replacement using a biodegradable porous polymerimplant. Knee Surgery Sports Traumatology Arthroscopy19 (3), 441–451.

Herwig, J., Egner, E., Buddecke, E., 1984. Chemical changes ofhuman knee joint menisci in various stages of degeneration.Annals of the Rheumatic Diseases 43 (4), 635–640.

Joshi, M.D., Suh, J.K., Marui, T., Woo, S.L., 1995. Interspeciesvariation of compressive biomechanical properties of themeniscus. Journal of Biomedical Materials Research 29 (7),823–828.

June, R.K., Ly, S., Fyhrie, D.P., 2009. Cartilage stress-relaxationproceeds slower at higher compressive strains. Archives ofBiochemistry and Biophysics 483 (1), 75–80.

Katsuragawa, Y., Saitoh, K., Tanaka, N., Wake, M., Ikeda, Y.,Furukawa, H., et al., 2010. Changes of human menisci in

osteoarthritic knee joints. Osteoarthritis and Cartilage 18 (9),1133–1143.

Kelm L.N. Hysteresis Response of Articular Cartilage to CyclicLoading. Rice University, Houston, TX, MSc thesis, 2002.

Kazemi, M., Li, L.P., Savard, P., Buschmann, M.D., 2011. Creepbehavior of the intact and meniscectomy knee joints. Journalof the Mechanical Behavior of Biomedical Materials 4 (7),1351–1358.

Khan, I., Smith, N., Jones, E., Finch, D.S., Cameron, R.E., 2005.Analysis and evaluation of a biomedical polycarbonateurethane tested in an in vitro study and an ovine arthroplastymodel. Part I: materials selection and evaluation. Biomaterials26, 621–631.

Kobayashi, M., Chang, Y.S., Oka, M., 2005. A two year in vivo studyof polyvinyl alcohol-hydrogel (PVA-H) artificial meniscus.Biomaterials 26 (16), 3243–3248.

Kurosawa, H., Fukubayashi, T., Nakajima, H., 1980. Load-BearingMode of the Knee Joint: Physical Behavior of the Knee JointWith or Without Menisci. Clinical Orthopaedics and RelatedResearch 149, 283–290.

Li, L.P., Buschmann, M.D., Shirazi-Adl, A., 2003. Strain-ratedependent stiffness of articular cartilage in unconfinedcompression. Journal of Biomechanical Engineering 125,161–168.

Linder-Ganz, E., Elsner, J.J., Danino, A., Guilak, F., Shterling, A.,2010. A novel quantitative approach for evaluating contactmechanics of meniscal replacements. Journal of BiomechalEngineering 132 (2), 024501.

Levy, I.M., Torzilli, P.A., Gould, J.D., Warren, R.F., 1989. The effectof lateral meniscectomy on motion of the knee. Journal ofBone and Joint Surgery: American Volume 71, 401–406.

Lohmander, L.S., Englund, P.M., Dahl, L.L., Roos, E.M., 2007.The long-term consequence of anterior cruciate ligament andmeniscus injuries: osteoarthritis. American Journal of SportsMedicine 35 (10), 1756–1769.

Mak, A.F., 1986. The apparent viscoelastic behavior of articularcartilage - the contributions from the intrinsic matrixviscoelasticity and interstitial fluid flows. Journal ofBiomechanical Engineering 108, 123–130.

McDermott, I.D., Amis, A.A., 2006. The consequences ofmeniscectomy. Journal of Bone and Joint Surgery: BritishVolume.88-B, 1549–1556.

McDermott, I.D., Masouros, S.D., Amis, A.A., 2008. Biomechanicsof the menisci of the knee. Current Orthopaedics. 22 (3),193–201.

McCrum, N.G., Buckley, C.P., Bucknall, C.B., 2003. Principles ofpolymer Engineering. Oxford University Press.

Natoli, R.M., Athanasiou, K.M., 2009. Traumatic loading ofarticular cartilage: mechanical and biological responses andpost-injury treatment. Biorheology 46, 451–485.

Oliveira, B.F., Creus, G.J., 2004. An analytical–numericalframework for the study of ageing in fibre reinforcedpolymer composites. Composite Structures 65 (3–4),443–457.

Park, S., Hung, C.T., Ateshian, G.A., 2004. Mechanical responseof bovine articular cartilage under dynamic unconfinedcompression loading at physiological stress levels.Osteoarthritis and Cartilage 12 (1), 65–73.

Rytter, S., Jensen, L.K., Bonde, J.P., Jurik, A.G., Egund, N., 2009.Occupational kneeling and meniscal tears: a magneticresonance imaging study in floor layers. Journal ofRheumatology 36 (7), 1512–1519.

Schatzmann, L., Brunner, P., Staubli, H.U., 1998. Effect of cyclicpreconditioning on the tensile properties of humanquadriceps tendons and patellar ligaments. Knee Surgery,Sports Traumatology, Arthroscopy 6 (1), S56–S61.

Shirazi, R., Shirazi-Adl, A., Hurtig, M., 2008. Role of cartilagecollagen fibrils networks in knee joint biomechanics

j o u r n a l o f t h e m e c h a n i c a l b e h a v i o r o f b i o m e d i c a l m a t e r i a l s 2 9 ( 2 0 1 4 ) 4 2 – 5 5 55

under compression. Journal of Biomechanics 41 (16),3340–3348.

Shelburne, K.B., Torry, M.R., Pandy, M.G., 2006. Contributions ofmuscles, ligaments, and the ground-reaction force totibiofemoral joint loading during normal gait. Journal ofOrthopedic Research 24 (10), 1983–1990.

Suh, J.-K., DiSilvestro, M.R., 1999. Biphasic poroviscoelasticbehavior of hydrated biological soft tissue. Journal ofBiomechanical Engineering 66, 528–535.

Terzidis, I.P., Christodoulou, A., Ploumis, A., Givissis, P.,Natsis, K., Koimtzis, M., 2006. Meniscal tear characteristicsin young athletes with a stable knee: arthroscopicevaluation. American Journal of Sports Medicine 34 (7),1170–1175.

Tissakht, M., Ahmed, A.M., 1995. Tensile stress-straincharacteristics of the human meniscal material. Journalof Biomechanics 28 (4), 411–422.

Uezaki, N., Kobayashi, A., Matsushige, K., 1979. The viscoelasticproperties of the human semilunar cartilage. Journalof Biomechanics 12 (1), 65–73.

van Tienen, T.G., Heijkants, R.G., Buma, P., de Groot, J.H.,Pennings, A.J., Veth, R.P., 2002. Tissue ingrowth anddegradation of two biodegradable porous polymers withdifferent porosities and pore sizes. Biomaterials 23 (8),1731–1738.

Winby C.R., Lloyd D.G., Besier T.F., Kirk T.B. 2009. Muscle andexternal load contribution to knee joint contact loads duringnormal gait. Journal of Biomechanics.16;42(14):2294–2300.

Ward, I.M., Sweeney, J., 2004. An Introduction to the MechanicalProperties of Solid Polymers, second ed. John Wiley & SonsLtd., Chichester, England.