Journal of Biomechanics 37 (2004) 1647–1653

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doi:10.1016/j.jb

Characterization of in vivo Achilles tendon forces in rabbits duringtreadmill locomotion at varying speeds and inclinations

John R. Westa, Natalia Juncosaa, Marc T. Gallowayb, Gregory P. Boivinc,David L. Butlera,*

aNoyes-Giannestras Biomechanics Laboratories, Department of Biomedical Engineering, University of Cincinnati, Cincinnati, OH 45221-0041, USAbCincinnati Sports Medicine and Orthopaedic Center, Cincinnati, Ohio, USA

cDepartment of Pathology and Laboratory Medicine, University of Cincinnati, USA

Accepted 18 February 2004

Abstract

The objective of this study was to test the hypothesis that increasing the speed and inclination of the treadmill increases the peak

Achilles tendon forces and their rates of rise and fall in force. Implantable force transducers (IFT) were inserted in the confluence of

the medial and lateral heads of the left gastrocnemius tendon in 11 rabbits. IFT voltages were successfully recorded in 8 animals as

the animals hopped on a treadmill at each of two speeds (0.1 and 0.3mph) and inclinations (0� and 12�). Instrumented tendons were

isolated shortly after sacrifice and calibrated. Contralateral unoperated tendons were failed in uniaxial tension to determine

maximum or failure force, from which safety factor (ratio of maximum force to peak in vivo force) was calculated for each activity.

Peak force and the rates of rise and fall in force significantly increased with increasing treadmill inclination ðpo0:001Þ: Safety factorsaveraged 30.877.5 for quiet standing, 7.072.9 for level hopping, and 5.270.7 for inclined hopping (mean7SEM). These in vivo

force parameters will help tissue engineers better design functional tissue engineered constructs for rabbit Achilles tendon and other

tendon repairs. Force patterns can also serve as input data for mechanical stimulation of tissue-engineered constructs in culture.

Such approaches are expected to help accelerate tendon repair after injury.

r 2004 Elsevier Ltd. All rights reserved.

Keywords: Functional tissue engineering; Achilles tendon; In vivo force; Implantable force transducer; Rabbit

1. Introduction

To better design load-bearing repair and graft tissuesafter tendon injury, it is imperative to understand theforce environment in which these tissues are expected toperform. While investigators have measured in vivoforces in the human Achilles tendon (Komi, 1990; Komiet al., 1992; Fukashiro et al., 1995), in several tendons inthe rabbit (Komi et al., 1996; Malaviya et al., 1998;Juncosa et al., 2003), cat (Ronsky et al., 1995), and goatmodels (Korvick et al., 1996) using a variety of devices,these data have not routinely been used as design criteriafor repair of these structures. Utilizing such a strategycan avoid failure of repair tissues after surgery andpotentially improve their performance during activitiesof daily living. This approach of designing repairs to

ing author. Tel.: 513-556-4167; fax: 513-556-4162.

ess: david.butler@uc.edu (D.L. Butler).

e front matter r 2004 Elsevier Ltd. All rights reserved.

iomech.2004.02.019

meet and even exceed expected in vivo forces is one ofthe principles of functional tissue engineering or FTE(Butler et al., 2000; Guilak et al., 2001, 2003)Although mesenchymal stem cells (MSCs) offer an

attractive treatment modality to improve tendon repair,this approach has still not incorporated tendon functionin its design or construct preparation. Inserting MSCs inrabbit patellar tendon window defects results in smallimprovements in rabbit patellar tendon repair biome-chanics (Awad et al., 1999). Organization of these cellsin the construct using a tensioned suture causes uniformorientation of the cells before surgery (Awad et al.,2000) and improves rabbit Achilles and patellar tendonrepair between 4 and 26 weeks after surgery (Younget al., 1998; Awad et al., 2003). However, a comparisonof the properties of cell-seeded repairs to properties ofoperated controls is not as valuable as a comparison tonormal tissue under normal operating forces. Specificcharacteristics of these in vivo mechanical signals such

ARTICLE IN PRESSJ.R. West et al. / Journal of Biomechanics 37 (2004) 1647–16531648

as peak forces, rates of rise and fall in tendon forces, andduty cycles might even permit repair biomechanics to beoptimized after surgery. These mechanical signals,which will likely be tissue and animal specific, are stillnot known in the rabbit Achilles tendon for our injurymodel.The long-term objectives of this research are to better

understand the in vivo force environments in varioustendon models and to develop design criteria forfunctional tissue engineered constructs containingMSCs. The specific purpose of this study was tocharacterize the detailed in vivo force patterns in therabbit Achilles tendon using implantable force transdu-cers (IFTs) (Xu et al., 1992; Glos et al., 1993) inserted inthe tissue midsubstance. We sought to test the hypoth-esis that increasing the speed and inclination of theactivity increases the peak Achilles tendon forces andtheir rates of rise and fall in force.

2. Methods

Eleven one-year old (weighing 4.770.1 kg; mean7SEM), female New Zealand White rabbits (MyrtlesRabbitry, Thompson Station, TN) were purchased forthe study. Data were acquired from eight animals (datafrom three animals were lost due to technical failure ofthe transducers). Skeletally mature rabbits were selectedto increase the likelihood of soft tissue failures duringtensile failure testing (Noyes and Grood, 1976).Tendon forces were measured using implantable force

transducers (Xu et al., 1992; Glos et al., 1993; Holdenet al., 1995; Korvick et al., 1996; Malaviya et al., 1998;Juncosa et al., 2003) (Fig. 1). Each IFT, with straingages in a half-bridge configuration, was inserted into apocket in the body of the tendon to avoid disruption bythe motion of surrounding tissues. The IFT producesvoltages that are nearly proportional to the tensile forcesin the tissue (Xu et al., 1992; Glos et al., 1993; Malaviyaet al., 1998; Juncosa et al., 2003).Five activity levels were evaluated in all animals.

Measurements were made during: quiet standing (QS),to simulate in vivo forces during disuse or inactivity;level hopping (LH) at 0.04m/s (0.1mph) and 0.13m/s(0.3mph), to simulate ‘‘in-cage’’ movements; and for

Fig. 1. Implantable force transducer (IFT) in tissue (Xu et al., 1992).

12� of inclined hopping (IH) at 0.04m/s (0.1mph) and0.13m/s (0.3mph), to simulate more vigorous exercise.Rabbits employed in this study were unable to hop onthe treadmill for more than 2min at a higher treadmillspeed than 0.13m/s. Animals were trained and exercisedon a treadmill as previously described (Juncosa et al.,2003). Preoperative training was conducted daily for 1week prior to IFT implantation (Oyen-Tiesma et al.,1998). Order of activity was randomized before testingusing a random number generator. After completion ofthe tests, the treadmill was then stopped at 0 degrees ofinclination so IFT signals could be re-recorded duringquiet standing.Animals were procured and surgically treated using

procedures approved by the Institutional Animal Careand Use Committee. Each rabbit was anesthetized by aninjection of ketamine hydrochloride (0.5ml/kg) andacepromazine (0.01ml/kg) in the psoas major muscle.Anesthesia was maintained during surgery using iso-flurane gas (1.5%–2.5%). The hind limb, lumbar regionand posterior neck of the rabbit were shaved andsterilely prepped. The left Achilles tendon was exposedthrough a lateral skin incision and the IFT was placed ina slit created in the confluence of the medial and lateralheads of the gastrocnemius tendons, 1 cm above thecalcaneus. The transducer was oriented such that itslong axis was parallel to the tendon fibers and its convexsurface faced posteriorly (Fig. 2). The tendon incisionwas closed with interrupted 5-0 Prolene sutures and theIFT lead wires were tethered to the knee’s lateral fasciafor strain relief and tunneled subcutaneously across theleft flank to the back of the neck where the electronicconnector was inaccessible to the rabbit. Skin incisionswere closed with 5-0 Prolene sutures. The rabbit wasthen given buprenorphine hydrochloride (0.05mg/kg),and its recovery was monitored every 20min untilit independently maintained sternal recumbency. AnElizabethan collar was placed around the neck toprotect the connector and lead wires from damage.

Fig. 2. Approximate placement of IFT in Achilles tendon between the

calcaneus and the gastrocnemius muscles.

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Fig. 3. A typical calibration curve is shown. Note the nearly linear

relationship between IFT voltage output and Achilles tendon force

ðr ¼ 0:98Þ:

Fig. 4. Uniaxial failure test using freeze clamps to secure the proximal

tendon and PMMA to secure the distal calcaneus. The freeze clamps

are attached to the moving actuator and the distal grip is attached to

the rigid load cell.

J.R. West et al. / Journal of Biomechanics 37 (2004) 1647–1653 1649

The suture site was evaluated daily to ensure there wereno complications.To verify equal force distribution between the hind

limbs, animals were placed on an electronic pressuremeasurement pad (Model 5250, Tekscan Inc., SouthBoston, MA) to measure ground reaction forces of bothlimbs during quiet standing. Forces transmitted by eachlimb were isolated and used to calculate operated-to-unoperated ground reaction force ratios.Data were collected 3 days after surgery at UC’s

Department of Laboratory Animal Medicine. Thisrecovery interval was selected to be long enough tolimit the effects of surgery on the movement of theanimal and short enough to avoid the risk of damage tothe IFT. Voltages were recorded using a laptopcomputer (Model Micron Transport Trek2, MicronPC,LLC, Nampa ID), signal conditioner-amplifier (Micro-strain, Burlington, VT), and data acquisition card (12bit National Instruments DAQ Card-500, NationalInstruments, Austin, TX) that were all directly inter-faced with the connector exiting the skin behindthe ears.Each animal was tested during each activity for ten

complete bouts, each bout consisting of three hops fromthe back to the front of the treadmill enclosure. Only thesecond hop was used to compute in vivo forces sincedata from the first hop could have been affected by therabbit accelerating from a complete stop at the back ofthe treadmill enclosure while data from the third hopcould have been influenced by the rabbit preparing tostop near the front of the enclosure. Each animal wasthen euthanized using sodium pentobarbital injection(100mg/kg).Each IFT/tendon complex was calibrated within 1 h

of testing and within 20min of euthanasia. The medialand lateral heads of the gastrocnemius muscle weretransected at the muscle-tendon junction and a #5Ethibond suture (Ethicon, Somerville, NJ) was securedto the Achilles tendon approximately 3 cm proximal tothe IFT implantation site. Using a calibrated springscale (Remington Arms, Madison, NC) increasing forcewas applied in 4.5N (1 lb) increments from 4.5N (1 lb)to 90N (20 lb) while simultaneously recording IFTvoltages (Fig. 3). After calibration, the rates of riseand fall were calculated using a linear curve fit betweenthe minimum and maximum force values for eachsecond hop as previously described (Juncosa et al.,2003).To determine safety factors for each of the three

activities, peak in vivo force and stress for each repairwere compared to the maximum force and maximumstress for the contralateral normal Achilles tendon,respectively. The contralateral tendon was used becausesurgery and the presence of the transducer could haveaffected the failure properties in the instrumentedtendon. Using previously described techniques (Butler

et al., 1984), the cross sectional area of the normaltendon was first measured at three locations andaveraged using a load-applied area micrometer at0.12MPa. The distal calcaneal bone block was em-bedded in a methylmethacrylate block (Dentsply Inc.,York, PA) and secured to the base plate on a materialtest machine (Model 8501, Instron Corp, Canton, MA).The proximal 2 cm of the tendon was connected to themachine’s actuator using freeze clamps (Fig. 4). Theinitial length was then measured between the grips(28.072.3mm) and the tissues were preconditioned intension for 50 cycles between 0 and 100N (approxi-mately 25% of failure force or 4% strain) at a constant

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Fig. 5. Typical in vivo force–time pattern transmitted by the rabbit

Achilles tendon during level hopping. Only the second hop is shown.

See text for definitions of each phase of the hop sequence.

J.R. West et al. / Journal of Biomechanics 37 (2004) 1647–16531650

strain rate of 10%/s. Each specimen was then failed intension at a rate of 8.5mm/s (B30%/s) while con-tinuously recording force and displacement. Safetyfactors for each activity were computed by dividingthe failure force or stress of the unoperated tendon bythe peak in vivo force or stress measured experimentallyin the contralateral tendon of each animal.The sample size ðn ¼ 8Þ was determined using a power

analysis to detect a 30% effect in peak force with apower of 85%. Analysis of Variance (ANOVA) wasconducted to determine the effects of treadmill speedand inclination on in vivo force response measures fromthe second hops of all activities. Similar to a previousreport (Juncosa et al., 2003), a randomized design wasemployed with Achilles tendon as the experimentalblock. A 2� 2 factorial, within-block treatment struc-ture was utilized with inclination (0� and 12�) and speed(0.04 and 0.13m/s) as the two factors. A mixed-modelanalysis was used with specimens as the random factor,all other factors being fixed. This design separated inter-animal variance from the measurement error variance.The ‘F’ statistic was obtained using p-values withBonferroni adjustment to account for multiple compar-isons among contrasts. All conclusions regarding effectof activity on in vivo force measurements were made atthe a ¼ 0:05 experiment-wise level. Foot contact forcemeasurements for the operated and unoperated limbswere evaluated using paired-t statistics.

Fig. 6. Treadmill inclination had a significant effect on peak tendon

forces ðpo0:001Þ: Peak forces (mean7SEM) are shown for different

levels of activity: LH/LS Level hopping at 0.04m/s, LH/HS Level

hopping at 0.13m/s, IH/LS Inclined hopping at 0.04m/s, IH/HS

Inclined hopping at 0.13m/s. HS: High Speed, LS: Low speed.

3. Results

Three transducers failed among the 11 animals tested.One IFT developed electronic noise of unknown origin,a second transducer failed at the junction of the leadwires with the transducer, and data from a third devicewere not utilized since the calculated in vivo forces fromcalibration were more than an order of magnitudegreater than for all other animals tested. The calibrationcurve relating voltage output and force input was nearlylinear up to 90N ðr ¼ 0:98Þ: No creep elongation wasvisually observed during the calibration procedurealthough elongation was not specifically measured.In vivo forces always remained greater than zero and

oscillated during hopping. Quiet standing (QS) forcesremained constant, averaging 16.372.2 N (mean7-SEM). Forces oscillated during hopping about a baseline

force that was above QS, typically dropping with anegative, almost linear initial slope to a minimum force

before increasing at a nearly constant rate of rise topeak force. The force then dropped again (rate of fall)before returning at a positive final slope to baseline force

(Fig. 5).Although increasing treadmill inclination significantly

increased peak force and the rates of rise and fallin force ðpo0:001Þ; increasing the level of treadmill

speed did not ðp ¼ 0:3Þ: Peak force was 57.770.5N forlevel hopping and 76.670.6N for inclined hopping(mean7SEM; Fig. 6). The rate of rise for combined 0.04and 0.13m/s speeds were 164.271.6N/s for levelhopping and 196.671.0N/s for inclined hopping(mean7SEM). The rate of fall for combined speedswas 163.171.5N/s for level hopping and 197.871.0N/sfor inclined hopping (mean7SEM). Force rise timeswere approximately 0.3 s regardless of activity level.Levels of treadmill inclination and speed did notsignificantly affect baseline force, initial slope, minimumforce, and final slope (Table 1, p ¼ 0:4).Safety factors decreased significantly with increasing

activity level from 30.877.5N during quiet standing to7.071.0 for level hopping to 5.270.7 for inclinedhopping (mean7SEM; Table 2). Maximum force forunoperated Achilles tendon specimens averaged402.6756.4N (mean7SEM; Fig. 7).Surgery did not significantly affect the distribution of

ground reaction forces between the operated andunoperated limbs during quiet standing ðp ¼ 0:9Þ: The

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Table 1

Activity level had no effect on baseline force, initial slope, minimum force, and final slope (p ¼ 0:4; mean7SEM)

0� 12�

0.1mph 0.3mph 0.1mph 0.3mph

Baseline force (N) 24.370.5 24.470.4 24.370.4 24.170.3

Initial slope (N/s) 52.670.8 52.370.7 53.870.5 53.870.4

Minimum force (N) 8.970.5 7.870.3 8.570.3 8.370.3

Final slope (N/s) 51.370.9 51.870.7 53.570.3 52.171.0

Table 2

Peak in vivo force, stress and strain during inclined hopping compared to ultimate force, stress and strain. Note that the peak in vivo force is

approximately 20% of maximum force

In vivo IH data Subject no. Weight (kg) Peak force (N) CSA (mm2) Peak stress (MPa) Peak strain (%)

1 4.6 78.6 12.6 6.2 4.4

2 4.4 74.9 9.5 7.9 4.1

3 4.5 78.7 9.7 8.1 4.5

4 4.6 75.5 10.5 7.2 3.5

5 4.8 76.2 14.5 5.3 3.9

6 5.2 79.5 13.4 5.9 3.7

7 4.7 78.9 11.3 7.0 4.1

8 5 80.2 13.1 6.1 3.8

Mean 4.7 77.8 11.8 6.7 4.0

SEM 0.3 0.7 0.6 0.3 0.1

Tensile failure data Subject no.a Weight (kg) Ult force (N) CSA (mm2) Ult stress (MPa) Ult strain (%)

1 4.6 403.3 12.8 31.6 13.4

2 4.4 307.0 9.3 32.9 18.2

3 4.5 285.9 9.9 28.7 15.5

4 4.6 146.8 10.7 13.8 14.5

5 4.8 523.0 14.3 36.7 16.1

6 5.2 377.3 13.6 27.7 17.8

7 4.7 536.1 11.5 46.6 16.1

8 5 641.2 13.3 48.1 16.6

Mean 4.7 402.6 11.9 33.3 16.0

SEM 0.3 56.4 0.6 3.9 0.6

In vivo/ultimate ratio 19.3% 20.1% 25.0%

aMeasurements performed in contralateral tendon.

Fig. 7. Normal Achilles tendon force–displacement curve showing the

peak forces for each level of activity. Safety factors decreased

significantly with increasing activity level from 30.877.5 N during

quiet standing (QS) to 7.071.0 for level hopping (LH) to 5.270.7 for

inclined hopping (IH) (mean7SEM).

J.R. West et al. / Journal of Biomechanics 37 (2004) 1647–1653 1651

ratio of contact forces for the operated vs. unoperatedlimbs was 1.0370.04 (mean7SEM, Fig. 8). There was alinear relationship between the contact forces of theoperated and unoperated limbs ðr ¼ 0:99Þ:

4. Discussion

To test our hypothesis that tendon forces increasewith activity level and to develop design criteria fortissue engineered implants, we directly measured in vivoforces in the rabbit Achilles tendon. The responsemeasures chosen for this study were similar to thoseused previously (Malaviya et al., 1998; Juncosa et al.,2003) and constitute discrete measurable factors thatcould be reproduced during mechanical stimulation oftissue engineered constructs to be later used at surgery.

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Fig. 8. The relationship between the ground reaction forces of the

operated and unoperated limbs in different animals ðn ¼ 8Þ was linearðr ¼ 0:99Þ: The ground reaction forces for both limbs were almost

identical.

J.R. West et al. / Journal of Biomechanics 37 (2004) 1647–16531652

The study is not without limitations. First, gaitpatterns examined only 3 days after IFT installationmay have been affected by the surgery, althoughabnormal hind limb motion was not observed. Earlypostoperative measurements insured that the IFTsremained functional and avoided any noticeable tissuereaction to the presence of the device. Second, noattempt was made to monitor knee and ankle jointflexion angles during the five activities and calibration.Thus, joint kinematics could not be correlated within vivo tendon force patterns. Third, no attempt wasmade to relate the measured AT forces to functionalaspects of the gait cycle, such as impact and push-offphases. This analysis will require better visualization ofthe limb during hopping or instrumentation of the footor tread surface. Such correlations in future studiesmight elucidate the mechanisms for the observedtemporal force patterns. Fourth, despite controllingthe speed of the treadmill, the rate was slow (to avoidanimal fatigue or injury) and generally not the speed atwhich the animals were moving. Treadmill speed wasvaried between only 0.04 and 0.13m/s but the rabbitswould typically hop from the back to the front of theenclosure in a three-hop sequence at a more rapid speedthan the treadmill tread. They would then rest as thetread carried them to the back of the enclosure wherethey would begin another three-hop sequence. Changesin treadmill speed only altered the rest time and not thehopping speed. Better control of the animal’s speed anda broader range of speeds might provide more valuableinformation regarding normal Achilles tendon forces.Fifth, the AT in vivo forces were measured for relativelylow speeds that the animals could tolerate. Futurestudies should try and impose higher demands on therabbits so as to establish lower bounds of safety factors.Sixth, a single calibration factor, determined duringpost-mortem calibration at a single midrange limbposition, may not have accounted for changes in IFTsensitivity with changes in knee and ankle joint flexionangles.

Activity level significantly affected the magnitude ofthe in vivo forces developed by the Achilles tendon.Peak tendon force increased with increasing treadmillinclination in a manner consistent with other in vivoreports for the human Achilles (Komi, 1990; Komi et al.,1992), goat patellar (Korvick et al., 1996), cat gastro-cnemius (Herzog et al., 1993; Prilutsky et al., 1994), andrabbit patellar and flexor tendon models (Juncosa et al.,2003; Malaviya et al., 1998). The fact that this patternwas not observed for the cat soleus and tibialis anteriortendons (Gregor et al., 1991; Herzog et al., 1993;Prilutsky et al., 1994) may well reflect differences inmuscle fiber type or action, but this factor was notexamined. Future studies might include both electro-myography (EMG) and kinematics to learn more aboutmuscle force generation characteristics and correlationwith limb position.Activity level was also positively correlated with the

rates of rise and fall in Achilles tendon force ðpo0:001Þ:These rates are in fact similar to those found for therabbit patellar tendon from a previous study (Juncosaet al., 2003) as are the times required to rise to peakforce and fall to baseline values (0.3 s for both LH andIH activities).The pattern of force vs. time for the Achilles tendon

during hopping was different from the patellar tendon.The Achilles tendon experienced an initial drop in forcenot observed in the patellar tendon (Juncosa et al.,2003). This drop may be due to relaxation of thegastrocnemius muscle as the ankle flexes during therecoil phase of each hop. Alternatively this drop maybe caused by the fact that the two-joint gastrocnemiusmuscle–tendon unit may have been affected by theindependent motions of both the knee and ankle joints.This study cannot distinguish between these proposedmechanisms since these joint motions were not mea-sured. Future research should seek to understand theseparate roles of knee and ankle joint motion onAchilles tendon forces.Even for inclined hopping, the Achilles tendon

developed in vivo forces and stresses that, on averagewere approximately 19% of the tendon’s maximumforce and stress values (Table 2). The correspondingsafety factors (5.2 for IH and 7.0 for LH) are larger thanthose measured previously with IFTs in the goat patellartendon (Korvick et al., 1996) but smaller than thosemeasured in the rabbit patellar tendon (Juncosa et al.,2003). The lower activities assessed in the current study(designed to insure the safety of the animal) mayrepresent an upper bound on safety factor for theAchilles tendon. Tendon forces should be measured athigher speeds in the rabbit and other models to betterunderstand the degree to which this muscle–tendon unitcan be challenged in vivo.The results of this study provide important informa-

tion for tissue engineers preparing biologic constructs

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for implantation in the rabbit Achilles tendon. Ideallyin vivo force patterns would be applied to cell-seededcollagen gels to stimulate the creation of organizedtissues. However, because the gels lack mechanicalintegrity, tissue engineers might wish to impose in vivostrain signals rather than in vivo stresses duringmechanical stimulation experiments. Although thesepeak strains, estimated to range between 1% and 4%in the rabbit Achilles tendon, are similar to thosemeasured in the rabbit patellar tendon (Juncosa et al.,2003), the actual strain patterns will differ due to themore complex nature of the Achilles tendon signal. Asmechanical properties of the stimulated constructimprove in culture or alternative scaffolds are used, itmay be possible to deliver simulated in vivo forces to theconstructs prior to implantation. The rabbit Achillestendon offers one model system for investigating varioustreatment approaches using principles of functionaltissue engineering.

Acknowledgements

This research was supported a grant from theNational Institutes of Health (AR 46574) given to theUniversity of Cincinnati and by Cincinnati SportsMedicine Research and Education Foundation. Theauthors wish to thank Dr. Martin Levy of the Universityof Cincinnati’s Department of Quantitative Analysis forhis assistance in statistical analysis, Lori Pitstick foranimal surgery assistance, and Dr. Kent N. Bachus atthe University of Utah’s Department of OrthopaedicSurgery for loan of the freeze clamps used in maximumstrength testing.

References

Awad, H.A., Butler, D.L., et al., 1999. Autologous mesenchymal

stem cell-mediated repair of tendon. Tissue Engineering 5 (3),

267–277.

Awad, H.A., Butler, D.L., et al., 2000. In vitro characterization of

mesenchymal stem cell-seeded collagen scaffolds for tendon repair:

effects of initial seeding density on contraction kinetics. Journal of

Biomedical Materials Research 51 (2), 233–240.

Awad, H.A., Boivin, G.P., et al., 2003. Repair of patellar tendon

injuries using a cell-collagen composite. Journal of Orthopaedic

Research 21 (3), 420–431.

Butler, D.L., Grood, E.S., et al., 1984. Effects of structure and strain

measurement technique on the material properties of young human

tendons and fascia. Journal of Biomechanics 17 (8), 579–596.

Butler, D.L., Goldstein, S.A., Guilak, F., 2000. Functional tissue

engineering: the role of biomechanics. Journal of Biomechanical

Engineering 122 (6), 570–575.

Fukashiro, S., Komi, P.V., Jarvinen, M., Miyashita, M., 1995. In vivo

Achilles tendon loading during jumping in humans. European

Journal of Applied Physiology and Occupational Physiology 71 (5),

453–458.

Glos, D.L., Butler, D.L., et al., 1993. In vitro evaluation of an

implantable force transducer (IFT) in a patellar tendon model.

Journal of Biomechanical Engineering 115 (4A), 335–343.

Gregor, R.J., Komi, P.V., et al., 1991. A comparison of the triceps

surae and residual muscle moments at the ankle during cycling.

Journal of Biomechanics 24 (5), 287–297.

Guilak, F., Butler, D.L., et al., 2001. Functional tissue engineering: the

role of biomechanics in articular cartilage repair. Clinical

Orthopaedics (391 Suppl), S295–305.

Guilak, F., Butler, D.L., et al., 2003. Functional Tissue Engineering.

Springer, New York.

Herzog, W., Stano, A., et al., 1993. Telemetry system to record force

and EMG from cat ankle extensor and tibialis anterior muscles.

Journal of Biomechanics 26 (12), 1463–1471.

Holden, J.P., Grood, E.S., et al., 1995. Factors affecting sensitivity of a

transducer for measuring anterior cruciate ligament force. Journal

of Biomechanics 28 (1), 99–102.

Juncosa, N., West, J.R., et al., 2003. In-vivo forces used to develop

design parameters for tissue engineered implants for rabbit patellar

tendon repair. Journal of Biomechanics 36 (4), 483–488.

Komi, P.V., 1990. Relevance of in vivo force measurements to human

biomechanics. Journal of Biomechanics 23 (Suppl 1), 23–34.

Komi, P.V., Fukashiro, S., et al., 1992. Biomechanical loading of

Achilles tendon during normal locomotion. Clinical Sports and

Medicine 11 (3), 521–531.

Komi, P.V., Belli, A., et al., 1996. Optic fibre as a transducer of

tendomuscular forces. European Journal of Applied Physiology 72

(3), 278–280.

Korvick, D.L., Cummings, J.F., et al., 1996. The use of an implantable

force transducer to measure patellar tendon forces in goats (see

comments). Journal of Biomechanics 29 (4), 557–561.

Malaviya, P., Butler, D.L., et al., 1998. In vivo tendon forces correlate

with activity level and remain bounded: evidence in a rabbit flexor

tendon model. Journal of Biomechanics 31 (11), 1043–1049.

Noyes, F.R., Grood, E.S., 1976. The strength of the anterior cruciate

ligament in humans and Rhesus monkeys. Journal of Bone Joint

Surgery (Am) 58 (8), 1074–1082.

Oyen-Tiesma, M., Atkinson, P.J., et al., 1998. A method for promoting

regular exercise in rabbits involved in orthopeadics research.

Contemporary Topics 37 (6), 77–80.

Prilutsky, B.I., Herzog, W., et al., 1994. Force-sharing between cat

soleus and gastrocnemius muscles during walking: explanations

based on electrical activity, properties, and kinematics. Journal of

Biomechanics 27 (10), 1223–1235.

Ronsky, J.L., Herzog, W., et al., 1995. In vivo quantification of the cat

patellofemoral joint contact stresses and areas. Journal of

Biomechanics 28 (8), 977–983.

Xu, W.S., Butler, D.L., et al., 1992. Theoretical analysis of an

implantable force transducer for tendon and ligament structures.

Journal of Biomechanical Engineering 114 (2), 170–177.

Young, R.G., Butler, D.L., et al., 1998. Use of mesenchymal stem cells

in a collagen matrix for Achilles tendon repair. Journal of

Orthopaedics Research 16 (4), 406–413.