Opi juoksemaan avojaloin

Avojaloin on juoksu on maailman vanhin liikuntamuoto. Se on 2 miljoonaa vuotta vanha, ja siihen pystyvät niin nuoret kuin vanhat, miehet kuin naiset. Lue lisää avojaloin juoksun biomekaniikkasta, ja siitä kuinka ihminen pystyy juoksemalla päihittämään jopa nelijalkaisia juoksuun erikoistuneita nisäkkäitä.

Monet kuitenkin pelkäävät avojaloin juoksemista tai kokevat sen vaikeaksi. Pelko on ymmärrettävää, sillä meidät on kantavaimennettujen ja pronaatiotuettujen kenkien avulla opetettu juoksemaan täysin vastoin luontoa.

Väärän tavan poisoppiminen ei kuitenkaan ole vaikeaa. Olemme kaikki saaneet oikean tekniikan syntymälahjana, eikä avojaloin juoksemiseen liity mitään tunnettuja riskejä.

Missä vika kun juokseminen sattuu?

Tiesin ammattini puolesta jo melko paljon juoksun biomekaniikasta, kun juoksin vielä kantavaimennetuilla ja pronaatiotuetuilla jalkineilla. Muutoksen laukaisi lopulta se, että liikunta ei enää tuntunut hauskalta jatkuvien kipujen kanssa. Päättelin, että joko ihmistä ei ole luotu liikkumaan, tai sitten minä olin tehnyt jotain väärin. Päätin uskoa jälkimmäiseen.

Heittäessäni vanhat juoksukengät nurkkaan kärsin akillesjänteen, jalkapöydän ja alaselän kivusta. Jalkapöydän kipu oli kinkkisin. Sen hoitaminen avojaloin juoksulla tuntui yhtä järkevältä kuin tenniskyynärpään hakkaaminen pesäpallomailalla.

Ensimmäisen kuukauden jälkeen mitään uusia vammoja ei ole ilmennyt ja vanhat vaivat ovat alkaneet helpottaa. Alussa se tuntui lähes ihmeeltä, vaikka taustalla onkin ihan tavallista biologiaa; kehomme tykkää liikunnasta. Pohkeissa ja akillesjänteessä toki esiintyi alussa arkuutta, mikä on luonnollista, kun keho saa uusia ärsykkeitä.

Kuuntele jalkojasi

Oikealla juoksutekniikalla on mahdollisuus päästä eroon vaivoista, mutta vain jos malttaa kuunnella omia jalkojaan. Avojaloin juoksussa ei päde sanonta no pain no gain. Päinvastoin, avojaloin juoksun ydin on keveys.

Tavoitteena on toki siirtää rajojaan kauemmaksi, mutta tehdä se hallitusti, omaa kehoaan kuunnellen.

Alussa et osaa mennä kevyesti. Silloin on mentävä vähemmän. On kuunneltava viestejä omista jaloista. Jos jaloissa on arkuutta, harrasta muuta liikuntaa. Juokset toivottavasti vielä kymmeniä vuosia, jolloin yhdellä päivällä ei ole mitään merkitystä.

Oikea tekniikka heti haltuun

Jos avojaloin juoksemiseen saisi antaa vain yhden ohjeen, se kuuluisi: Laskeudu päkiälle. Mutta vaikka tämä on tärkein ohje, pelkästään sen avulla voi olla vaikeaa päästä eroon kantavaimennetuilla kengillä opitusta väärästä tekniikasta. Siksi kehotan kaikkia avojaloin juoksemista suunnittelevia tutustumaan alta löytyviin ohjeisiin, videoon ja kuviin ennen lenkille lähtöä.

Video 1. Amerikkalainen lääkäri Mark Cucuzzella opastaa avojaloin juoksun tekniikassa.

Kuva 1. New York Times Magazine julkaisi hiljattain artikkelin otsikolla “The Once and Future Way to Run”, joka sisältää mm. hyviä ohjeita avojaloin juoksun oppimiseen. Viereinen kuvasarja on kyseisestä artikkelista ja kuvatekstien suomennokset ovat alla. (Linkki alkuperäiseen artikkeliin, vaatii rekisteröitymisen.)

  1. Vie polvi ja vastakkainen käsi eteen. Pidä selkä suorana.
  2. Kuvittele astuvasi kaatuneen puunrungon yli; mitä kovempi vauhti, sitä paksumpi runko.
  3. Nilkan ei koskaan pitäisi ojentua polven eteen.
  4. Polven liike syntyy etureiden ja pakaran lihasten aktivaatiosta. Pidä muu jalka rentona.
  5. Valmista jalkaterä sulavaa kontaktia varten. Päkiä osuu maahan ensimmäisenä.
  6. Kantapää ainoastaan hipaisee maata. Kontaktin tapahduttua aloita kierros toisella jalalla.

Edellä mainittuja ohjeita ei tarvitse ajatella lopun elämäänsä. Väärän tekniikan poisoppiminen sujuu nopeasti. Tiedät juoksevasi oikein, kun et haluaisi lopettaa juoksemista lainkaan. Huomaat myös, että mm. talven liukkailla juoksu muuttuu helpoksi ja juoksumatolla juokseminenkaan ei enää riko jalkapohjia.

Harhaluuloja ja opittuja virheitä

Jokainen lapsi osaa juosta. Onhan juoksemisen ohjeet syntymässä sisällytetty keskushermostoomme. Matkalla aikuiseksi monet meistä unohtavat tämän taidon. Syynä voivat olla matkan varrella käyttämämme vääränlaiset jalkineet, tai halumme ymmärtää ja hallita niinkin luonnollista toimintaa kuin juoksemista.

Esimerkiksi: Monet ajattelevat, että juoksussa säästää energiaa, jos liikkuu mahdollisimman vähän pystysuunnassa. Tämä on väärin. Energiaa kuluu lähinnä silloin, kun kannatellaan kehon massaa eli kontaktivaiheessa. Niinpä juoksuun kuuluu lentovaihe. Ilman pystysunnassa tapahtuvaa liikettä taas ei olisi lentovaihetta. Luonto antoi meille joustavat jalat. Käytä niitä ja unohda laahaus.

Tyypillisiä kengän kanssa opittuja virheitä voivat olla:

  1. Jalka laahaa maata ja sitä kurotetaan pitkälle eteen. Laskeutuminen tapahtuu kaukana painopisteestä ja johtaa jarrutukseen
  2. Laskeudutaan kantapäälle, jolloin jalan luonnollinen iskunvaimennus ja energian varastointi ei toimi.

Jotkut myös ponnistavat voimakkaasti irti maasta. Tämä on täysin tarpeetonta. Kun jalka osuu maahan, se tietää mitä tehdä. 60% pohkeen työstä tekee akillesjänne ja sen jäykkyys on optimoitu juuri juoksua ajatellen. Siihen varastoitu energia antaa aina sopivan työntövoiman; mitä kovempi vauhti sitä suurempi energia.

Ovatko kaikki kengät pahasta?

Eivät. Tärkeintä on, ettei kengässä ole akillesjänteen venymisen estävää kantavaimennusta eikä jalkaterän normaalia muodonmuutosta estäviä tukirakenteita. On mm. täysin luonnollista, että kontaktin aikana jalkaterä levenee päkiän kohdalta, jolloin kapea kenkä voi olla kiusallisten vaivojen syy.

On vaikeaa löytää uusia juoksukenkiä, joissa kantapäätä ei olisi korotettu. Paras apu voikin löytyä ullakolta, johon on säilötty vanhat juoksukengät ajalta, jolloin urheilujalkineet olivat vain tasainen pohja ja siihen ommeltu kangas. Voit myös kokeilla leikata korotettujen juoksutossujen kantapäästä puukolla ylimääräisen 1-2 cm:ä pois. Oma yritykseni ei tuottanut toivottua tulosta, sillä operaation jälkeen kaarituki tökki entistä kivuliaammin vastaan, ja kengät joutivat lopulta roskakoriin.

Maailman parhaat juoksukengät ovat sandaalit. Niissä ei ole mitään jalan muodonmuutosta estäviä rakenteita, ja ne taatusti hengittävät. Ennen kuin tavallisista rantasandaaleista tulee juoksusandaalit, voit joutua tuunaamaan niitä ainakin nauhojen osalta niin, että nauhat kulkevat myös kantapään ympäri ja estävät sandaalien irtoamisen jalasta. Ohjeita omien juoksusandaalien valmistamiseen löytyy helposti googlettamalla esim. sanoilla running sandals tai huaraches.

Kuva 2. Itse tehdyt huarache juoksusandaalit. Nimi tulee Tarahumara intiaaneilta, jotka valmistavat juoksusandaaleja käytetyistä autonrenkaista. Itse hankin pohjamateriaalin ja nahkanauhat suutarilta.

Lopuksi se meidän suomessa asuvien juoksijoiden suurin kiusa: talvi. Sen varalle en osaa vielä antaa hyvää ja toimivaksi todettua vinkkiä. Ehdottomasti mielenkiintoisin lukemani kuului kuitenkin näin: Talvella juokseminen onnistuu, kun laittaa päällekkäin muutaman villasukan.

Mukavia kilometrejä luonnollisella askeleella.

Avojaloin juoksun biomekaniikka

Olemmeko syntyneet juoksemaan avojaloin?

Ihminen on harrastanut kestävyysjuoksua arviolta n. 2 miljoonaa vuotta – ja pärjännyt melko hyvin. Evoluutiolla on siis ollut hyvin aikaa tehdä kehosta juoksukelpoinen. Silti juoksukenkien valmistajat totesivat 70-luvulla, että todennäköisesti luonto teki asian väärin. Ja korjasivat sen. Vai korjasivatko he vain hyväuskoisten kuluttajien rahat?

Kuinka hyviä juoksijoita olemme?

Eläimiin verrattuna ihmistä pidetään huonona juoksijana. Tämä on harhaluulo, joka on todennäköisesti syntynyt siitä tosiasiasta, että moniin eläimiin verrattuna ihminen on heikko ja siksi surkea pikajuoksija. Lähin sukulaisemme, simpanssi, on ihmistä voimakkaampi. Paljain nyrkein tapahtuva aikuisen miehen ja simpanssiuroksen välinen yhteenotto päättyisi ihmisen kannalta onnettomasti, huolimatta simpanssiuroksen pienemmästä koosta. Mutta jos miehelle annettaisiin 50 metrin etumatka, ei simpanssi saisi miestä koskaan kiinni.

Ei ole kovin yllättävää, että ihminen pärjää kestävyysjuoksussa simpanssille, jolla on raskas ylävartalo ja lyhyet jalat. Sen sijaan yllättävää on, että ihminen pärjää pitkän matkan juoksussa monille nelijalkaisille juoksuun erikoistuneille elämille, kuten susille, aavikkokoirille, gnuuantiloopeille ja jopa hevosille. Vuosittain sadat tuhannet ihmiset selvittävät 42 km:n maratonin. Sudet ja muut laumassa saalistavat lihansyöjätkään eivät juokse tällaisia matkoja luonnossa – niiden tyypillinen päivätaival on alle 15 km:ä. Jopa hevoset hidastavat vauhtiaan pitkällä yli 20 km:n taipaleella, jolloin hyväkuntoinen juoksija voi juosta niitä nopeammin.

Savannilla saaliin sai kiinni juoksemalla

Mistä sitten johtuu, että kestävyysjuoksukykymme on kerrassaan mainio? Ihmisen vahvuuksinahan on pääasiassa pidetty terävää älyä ja näppäriä käsiä. Vastausta kysymykseen on etsittävä kaukaa ajoilta jolloin ihmisen piti selviytyä eloonjäämistaistelussa savannin muita petoja vastaan. Petojen aseina olivat nopeus, terävät kynnet ja hampaat saaliin tappamiseksi sekä erinomainen hämäränäkö. Ihmisellä oli hyvä värinäkö, tehokas jäähdytysjärjestelmä ja vapaat kädet joissa kantaa vettä ja aseita. Ihminen tarvitsi kaikkia näitä, kun se saalisti keskipäivän auringossa juoksemalla saaliinsa uuvuksiin (video 1).

Kuinka on mahdollista, että ihminen juoksee kiinni antiloopin, jonka huippunopeus on paljon ihmistä suurempi? Ihmisellä on yksi merkittävä etu muihin nisäkkäisiin nähden, alaston iho. Nisäkkäät poistavat lämpöä läähättämällä; kierrättämällä ilmaa hyvin tiiviillä tahdilla kurkun yläosassa, eikä lainkaan keuhkoissa. Koska läähätys ei ole hengittämistä, eläimen on pysähdyttävä läähättämään. Ihminen ei kärsi tästä rajoituksesta, vaan poistaa lämpöä tehokkaasti hikoilemalla runsaiden hikirauhasten ja olemattoman turkin ansioista. Jos metsästäjä ei hukkaa saaliin jälkiä, se saavuttaa antiloopin siinä vaiheessa, kun eläin uupuu ja pysähtyy. Pedot 1 – ihminen 1. Molemmat saivat saaliinsa.

Ottaen huomioon nämä ihmisen eläinkunnankin mittareilla merkittävät saavutukset kestävyysjuoksussa, on mielenkiintoista kysyä: Voiko jalan rakenne olla niin juoksuun soveltumaton, että sitä on korjattava urheilujalkineilla?

Video 1. BBC:n luontodokumentti havainnollistaa kuinka ihmiset pyydystivät saaliinsa hellittämättömällä takaa-ajolla (persistence hunting).

Juoksu ei ole nopeaa kävelyä

Kävely ja juoksu kuljettavat meitä eteenpäin, mutta siihen niiden yhtäläisyydet loppuvatkin. Biomekaanisesti kävely ja juoksu nimittäin eroavat merkittävästi toisistaan. Kävely perustuu heilurimekanismiin kun taas juoksu on kuin kumipallolla pomppimista. Kävelyssä painopiste siirtyy eteenpäin vuorottelevan tukijalan varassa samaan aikaan, kun vastakkainen jalka heilahtaa kuin seinäkellon heiluri. Tämä on kaksijalkaisten taloudellisin etenemistapa, mutta vain hyvin kapealla nopeusalueella, joka on keskimäärin noin 5 km/h. Optimimaalisen nopeusalueen ulkopuolella kävelyn taloudellisuus heikkenee jyrkästi. Juoksussa jalat toimivat kumipallon tavoin ohjaten törmäysenergian elastisiin tukirakenteisiin. Parhaimmillaan kävely on juoksua taloudellisempaa, mutta ei paljon. Kävelykin kuluttaa n. 2/3 juoksun energiasta kilometriä kohti. Juoksun merkittävin etu kävelyyn nähden on, että juoksun energiankulutus kilometriä kohden on nopeudesta riippumaton. Juostessa nopeammin kuluu siis luonnollisesti enemmän energiaa samassa ajassa, mutta ei samalla matkalla.

Kävely on hitaan vauhtinsa vuoksi turvallista, eikä vaadi pitkälle kehittyneitä suojamekanismeja, koska törmäysenergiat ovat pieniä. Juoksussa tilanne muuttuu. Vauhti ja törmäysenergiat kasvavat ja kantapään törmäys tuottaa suuren voimapiikin, joka välittyy luiden kautta aina selkärankaan ja lopulta päähän saakka. Toistuessaan jatkuvasti tämä voi aiheuttaa kudosvaurioita ja äärimmillään johtaa pysyvään juoksukyvyttömyyteen. Törmäyksiä vaimentamaan kenkiin kehitettiin kantapatjat. Ongelma ratkaistu? Ei ihan, sillä ensiksi kantapatjat eivät toimi ja toiseksi törmääminen kantapää edellä estää nilkan elastisten rakenteiden luonnollisen toiminnan.

Muodonmuutos kuuluu juoksuaskeleeseen

Jalan rakenne on erityisen hyvin sopeutunut juoksemiseen. Se sisältää runsaasti elastisia rakenteita, jotka voivat muuttaa hetkellisesti muotoaan jännitysvoimien alla ja palautua entiselleen. Kutsumme näitä elastisia rakenteita jänteiksi, ja niitä sijaitsee jalkaholvissa, nilkkanivelessä ja akillesjänteessä. Jänteiden tehtävänä on toimia pienten jousten tavoin, joille törmäysenergia hallitusti ohjataan askeleen tukivaiheen ensimmäisen puolikkaan aikana. Mekanismi suojaa kehoa vaurioilta. Parasta on, että törmäysenergiaa ei ainoastaan varastoida, vaan sitä myös vapautetaan tukivaiheen toisen puolikkaan aikana. Juoksu on taloudellisimmillaan silloin, kun kontakti tapahtuu päkiällä ja jalan normaali muodonmuutos sallitaan (video 2).

Kaikki liikkeet, jotka mahdollistavat jänteiden normaalin toiminnan, on estetty monissa moderneissa juoksujalkineissa: 1) Jalkaholvin muodonmuutoksen estää tukikaari, 2) Nilkan pronaation pronaatiotuki ja 3) Akillesjänteen venymisen kantapatja, joka ohjaa törmäyksen virheellisesti kantapäälle. Poikkeuksen tekevät ainoastaan ammattilaisille suunnatut kevyet ja vaimentamattomat juoksukengät. Kuitenkin, miksei tavallinen harrastaja saisi hyödyntää niitä rakenteita, jotka miljoonien vuosien evoluution seurauksena ovat taanneet meille juoksukyvyn, joka kestää vertailussa mitä tahansa eläinkunnan lajia vastaan?

Video 2. Professori Daniel E. Lieberman Harvardin yliopistosta kertoo kuinka avojaloin juoksu eroaa jalkinejuoksusta (www.nature.com/nature).

Avojaloin pääsee kevyemmin – ja turvallisesti

Ei ole vaikea keksiä ominaisuutta, joka tekee mistä tahansa juoksujalkineesta hyvän suojavarusteen. Jokainen joka on viiltänyt jalkapohjansa pullonsirpaleeseen allekirjoittaa pohjan merkityksen. Pohja onkin juoksujalkineen ainoa ominaisuus, josta on kuviteltavissa olevaa suojaa loukkaantumisia vastaan. Toisaalta avojaloin juoksevat vakuuttavat, että ongelmaa ei ole, sillä eihän kaikkein vaikeimpiin paikkoihin tarvitse jalkaansa työntää.

Tosiasiassa ei ole olemassa mitään tieteellistä näyttöä modernin juoksujalkineen urheiluvammoilta suojaavasta vaikutuksesta. Varmasti tiedämme ainoastaan että mitä kevyempi jalkine, sitä taloudellisempi. Suurten massojen liikuttelu jalkaterissä nimittäin vie paljon energiaa. Paljain jaloin juokseminen on energiataloudellisesti kannattavinta!

Ihmiskeho, jalka mukaan lukien, on kehittynyt niin hyvin juoksuun soveltuvaksi, että olemme pystyneet kestävyysjuoksussa päihittämään nelijalkaisia juoksuun erikoistuneita nisäkkäitä. Tämä on melko hämmästyttävä saavutus ja sen rinnalla ei tunnu uskottavalta, että saavutuksen mahdollistavat piirteet olisivat kehittyneet turhaan. Nilkka ja jalkaterä ovat tärkeitä iskunvaimentimia ja energian varastoijia. Jalkaholvi ja akillesjänne yhdessä kykenevät varastoimaan karkeasti puolet juoksuaskeleen törmäysenergiasta. Ei ole ainuttakaan syytä uskoa, että terve jalka tarvitsi juoksussa mitään ylimääräistä tukea tai suojavarustetta.

Toki jokainen jalka on yksilöllinen, ja on mahdollista, että erikoistapauksissa lisätuesta on hyötyä. Silti, moni täysin tervejalkainen saattaa erehtyä luulemaan, että hänen jaloissaan olisi jotain vikaa, vaikka vika itse asiassa on siinä, että jalkoja ei vain ole käytetty siihen tarkoitukseen, mihin ne ovat 2 miljoonan vuoden aikana kehittyneet.

Can human tendon fatigue in exercise?

Published in: Journal of Experimental Biology 213, 1259-1265 (2010)
First published online March 26, 2010 [Link to journal abstract]
doi: 10.1242/jeb.033514

In vivo mechanical response of human Achilles tendon to a single bout of hopping exercise

Jussi Peltonen1, Neil J. Cronin1,2, Janne Avela1, Taija Finni1.

1Neuromuscular Research Center, Department of Biology of Physical Activity, University of Jyväskylä, Finland

2Center for Sensory-Motor Interaction, Department of Health Science and Technology, Aalborg University, Denmark

Key words: Achilles tendon, stiffness, fatigue, ultrasonography

Abstract

Stiffness of the human Achilles tendon (AT) was determined in vivo before and after a single bout of hopping exercise. It was hypothesized, based on published data using in vitro specimens, that a reduction in AT stiffness may occur after just 1000 loading cycles at physiological stress levels. Ten healthy subjects performed two-legged hopping exercise consisting of 1150-2600 high impacts. Tendon stiffness was determined in several isometric ramp contractions [20%, 40%, 60%, 80% and 100% maximum voluntary contraction (MVC)] during which tendon elongation was measured using ultrasonography and two cameras. Tendon force was calculated by dividing measured ankle torque by magnetic resonance image-derived AT lever arm length. Tendon stiffness remained unchanged, being 430±200 N mm–1 before and 390±190 N mm–1 after the exercise (not significant). Despite the lack of changes in stiffness, maximum tendon force during MVC was reduced from 3.5±0.6 kN to 2.8±0.7kN (P<0.01). As the proposed decline in stiffness was not observed, it is concluded that mechanical fatigue did not take place in the AT of healthy individuals after a single bout of high impact exercise performed until exhaustion.

Introduction

Both animal and human tendons have been shown to fatigue in vitro under continuous static or cyclic loading. In vitro, fatigue is often expressed as a decrease in the ultimate stress of a tendon; the stress that the tendon can bear without rupture. When rupture occurs after static loading, it is referred to as a creep rupture, and when it occurs after cyclic loading, it is known as a fatigue rupture. Creep or fatigue ruptures have been demonstrated in wallaby tendons (Wang and Ker, 1995a; Wang et al. 1995b, Ker et al., 2000) and in sheep tendons (Pike et al., 2000), fatigue ruptures in human extensor digitorum longus (EDL) tendons (Schechtman and Bader, 1994; 1997; 2002), and creep and fatigue ruptures in human Achilles tendons (AT) (Wren et al., 2003). However, ultimate stress is not a valid parameter to describe tendon fatigue in living tissue. Fortunately, most authors have concluded that the slope of the linear region of the stress-strain curve, known as Young’s modulus, is an excellent indicator of tendon fatigue. At the instant of rupture, Young’s modulus has been shown to decrease to approximately half of its original value (Schechtman and Bader, 2002; Wren et al., 2003).

Instead of Young’s modulus, tendon stiffness is often used in in vivo studies to quantify tendon fatigue. The difference between stiffness and Young’s modulus is that tendon stiffness depends on both material properties and dimensions of the tendon, whereas Young’s modulus is determined by material properties alone. In terms of stiffness, if fatigue is induced, tendon stiffness decreases, and inversely, tendon compliance increases.

An increase in tendon and aponeurosis compliance in vivo was first reported by Kubo and colleagues (2001a, 2001b, 2005) in vastus lateralis (VL) muscle after isometric contractions. By contrast, no changes in compliance have been observed after dynamic contractions in VL tendon (Kubo et al., 2001b; Kubo et al., 2005; Mademli et al., 2008; Ullrich et al., 2007) or in gastrocnemius medialis (GAM) tendon (Mademli et al., 2006). These findings suggest that the type of contraction or strain rate would determine the severity of tendon fatigue. The effect of strain rate on tendon fatigue resistance has not been investigated in vivo but the mode of failure in knee ligaments of primate specimens changed from a predominance of avulsion fracture at the slow rate to ligament rupture at the fast rate (Noyes et al., 1974). This indicates that high strain rates increase the likelihood of tendon rupture. Furthermore, if exercise is repeated without sufficient recovery, fatigue effects may accumulate and eventually lead to tendon overuse injuries (Kannus and Josza, 1991; Galloway et al., 1992). Therefore, this study was designed to investigate tendon mechanical responses to a single bout of exercise, and consequently, hopefully, to help to prevent tendon injuries.

In the literature, the effects of dynamic exercise on tendon properties have only been examined after isolated eccentric and concentric contractions. Thus, there is a need for a study utilizing a large number of impacts at high strain rates that occur during natural locomotion in sports and physical activities. The aim of this study was to investigate human AT fatigue by calculating tendon stiffness before and after a single bout of hopping exercise. Based on data from human specimens, it was hypothesized that Achilles tendon fatigue may be induced after as little as 1000 loading cycles (Wren et al., 2003). Tendon stiffness was chosen as an indicator of tendon fatigue since its derivative, Young’s modulus, has been shown to be a valid indicator of tendon fatigue in specimen studies (Schechtman and Bader, 1994; 2002; Wang and Ker, 1995a; Wang et al., 1995b; Wren et al., 2003).

Materials and methods

Six males and four females (N=10) participated in this study. Their mean body mass, height and age were 65±10 kg, 171±8 cm and 26±3 years (table 1), respectively. All participants were physically active and had diverse sport backgrounds. Participants were informed about the procedures, benefits and possible risks involved in the study, and they signed a written consent. All methods were approved by the ethical committee of the University of Jyväskylä. The study conformed to the standards set by the Declaration of Helsinki.

Protocol overview

Illustration of the measurement setup in the ankle dynamometer. Subject was tightly anchored between the pedal and back support. Two cameras (side and ground) were set up to record heel and ultrasound probe displacement. The side camera’s optical axis was perpendicular to the figure plane and therefore the camera is not visible in the figure.

Data were acquired in an ankle dynamometer as illustrated in figure 1A. To measure torque, the right foot was firmly attached to a pedal with a force transducer (Precision TB5-C1, Raute, Nastola, Finland) installed at a constant distance from the pedal rotation axis. Torque was sampled with a 16-bit AD-board (Power 1401, CED Limited, Cambridge, England) at 1 kHz. Seat position was adjusted individually to obtain the following joint angles: ankle 90±3°, knee 180±3°and hip 120±3°. The axis of rotation of the ankle joint was carefully aligned with the rotational axis of the pedal. Two-dimensional displacement of the myotendinous junction (MTJ) of GAM was recorded with a sampling frequency of 50Hz using a 6 cm linear array ultrasound (US) probe (UST-5712, Aloka, Tokyo, Japan) and imaging unit (Prosound Alpha 10, Aloka, Tokyo, Japan). US images were captured on VHS tape (576i at 50Hz). A similar procedure has frequently been used in studies of muscle mechanics during physical activities (Arampatzis et al., 2006; Lichtwark et al., 2007). The probe was placed 2 cm medial to the position where medial and lateral gastrocnemius muscles join. The probe was supported by a custom-made cast, and secured by elastic bandages wrapped around the leg to prevent probe movement during exercise. The lack of movement of the probe relative to skin was verified by marking the location of the probe on the skin prior to exercise. Movement of the US probe and heel during isometric contractions was controlled by setting two video cameras (Sony HDR-HC3, Tokyo, Japan), 576i at 50Hz) perpendicular to each other. The ground camera captured movement of a heel marker on the posterior surface of the calcaneus, and the side camera captured the coordinates of US probe markers. An EMG electrode pair (Ag/AgCl, 13.2 mm2, Blue Sensor M, Ambu A/S, Ballerup, Denmark) was placed over tibialis anterior (TA) to quantify antagonist activity during plantarflexions. The location of the electrode pair (inter-electrode distance 2 cm) was between the muscle mid belly and distal tendon. EMG signals were amplified (gain: 1000) and band-pass filtered (range: 10-1000 Hz) before being fed to the AD-board (Power 1401) for sampling at 1 kHz. All signals were synchronized by digital pulses.

Timeline of the current protocol. Exercise duration varied between subjects from 8 to 22 minutes. MVC, maximum voluntary contraction.

The timeline and content of the experimental protocol are shown in figure 1B. Experiments started with tendon pre-conditioning, which is a phenomenon that is characterized by increased stretching of tendon under constant repeated loading in animals (Abrahams, 1967) and humans (Maganaris and Paul, 2002; Maganaris, 2003). To distinguish short-term tendon conditioning from long-term tendon fatigue, five isometric contractions up to 80% of maximum were performed before the measurements, because this has been shown to be adequate to remove the conditioning effect (Maganaris, 2003). Isometric maximum voluntary contraction (MVC) force was measured as the best out of three trials. Tendon stiffness was then determined pre- and post-exercise using force and length data from the best MVC and five isometric ramp contractions up to the levels of 80%, 60%, 40%, 20% and 100%, measured at this order. The time between two consecutive isometric contractions was approx. ~1 min and during contraction, isometric target force was maintained for 2 seconds to stabilize the tendon deformation (figure 2). It was assumed that the effects of pre-conditioning were maintained throughout the experiment. The exact time course of recovery from pre-conditioning is not clear in vivo, but in vitro, Viidik (1973) reported that when static state was achieved, no elastic recovery was observed during the following 240 minutes in resting connective tissue.

Example of a single 80% ramp contraction (before exercise). Calculated tendon force (solid line) and measured elongation (broken line) are drawn over time. Achilles tendon force (∆F) and tendon elongation (∆L) were calculated as the difference between initial and maximum values that were taken as the mean over the gray shaded area.

MVC was measured independently before and after the exercise and all ramp contraction target levels (%) were calculated accordingly. A practical consequence of this is that tendon stiffness was determined at a lower absolute force level after the exercise. As demonstrated by in vivo stress-strain curves (Lichtwark and Wilson, 2005), stiffness is independent of the force in the linear region. Therefore a constant effort level was preferred over a constant force level (which could not have been guaranteed anyway because tendon force was calculated and not measured).

Two-legged hopping (feet together) on a contact mat (custom-made, capacitor switch) was performed until exhaustion. Loading was focused on the triceps surae muscles. Subjects were instructed to hop at their preferred frequency (range: 2.0-2.5 Hz) and strongly activate their calf muscles to keep the heels off the ground during the contact phase. During exercise, flight time ranged from 0.2 to 0.4 seconds (coinciding with a jump height of 5-20 cm) and the total number of jumps ranged between 1150 and 2600 (table 1).

Magnetic resonance images

(A) Achilles tendon cross-sectional area and (B) lever arm length were calculated from magnetic resonance images. The lever arm length was defined as the perpendicular distance from the midpoint of the superior surface of the trochlea of the talus to the line of action of Achilles tendon. Ankle angle was 90 degrees (sole of the foot perpendicular to shank).

Magnetic resonance images (MRI) of the lower right leg were taken one week before the exercise. Lever arm length and tendon cross-sectional area (CSA) were analyzed from MR images, which were acquired in axial (repetition time = 4500.00, echo time = 10.90, flip angle = 90.00, matrix = 256×256 and slice thickness = 7.00 mm) and sagittal orientations (repetition time = 100.00, echo time = 3.70, flip angle = 60.00, matrix = 256×256 and slice thickness = 5.00 mm). CSA was determined by manually outlining the AT in axial images along the tendon from the calcaneus to the MTJ of GAM. The smallest CSA (located 1.7-5.1 cm from the calcaneus) was used for later calculations (Kongsgaard et al., 2005). The smallest CSA has the greatest stress concentration, and failures usually begin there during cyclic fatigue tests (Wren et al., 2003). From sagittal images, AT lever arm length was determined as the perpendicular distance from the midpoint of the superior surface of the trochlea of the talus to the line of action of AT, a method that is similar to that previously used by Maganaris (2001) and Finni et al. (2006). Ankle angle was 90 degrees (sole of the foot perpendicular to shank) (Figure 3).

Data analysis and processing

AT force (∆F) was calculated by dividing measured ankle torque by MRI derived AT lever arm length. To yield ∆F, initial tendon force before contraction was subtracted from maximum tendon force at the end of the plateau phase of each isometric contraction as shown in figure 2. It was assumed that all plantar flexion force was transmitted through the AT and that the lever arm length remained constant. It is known that AT moment arm does not necessarily remain constant during muscle activity but variations are mainly caused by large joint rotations (Rugg et al., 1990; Maganaris et al., 2000), which were not observed here (< 5°), and therefore a constant moment arm is often used (e.g. Fukunaga et al., 1996). A TA activity correction method (Mademli et al., 2004) was applied to the measured force values if necessary. If TA EMG activity was observed during isometric plantarflexion, a counter-movement (dorsiflexion) was performed to produce similar TA EMG activity. Dorsiflexion torque was then added to the plantarflexion torque to estimate true AT force. Seven out of ten subjects produced ankle plantar flexion without TA activity and for the remaining three subjects, the correction was always less than 5%. Tendon stress was calculated by dividing tendon force by the smallest tendon CSA deduced from MRI.

Schematic representation of two coordinate systems. Achilles tendon length was determined as the distance between insertion (A) and myotendinous junction (B). Coordinates of myotendinous junction were transformed from the probe coordinate system to the laboratory coordinate system. Note that the vertical axis of the ultrasound (US) image is flipped to match image orientation in the US imaging unit. SOL = soleus, MG = medial gastrocnemius.

Tendon elongation (∆L) was also determined between initial and maximum tendon length during each isometric contraction and the average of the initial tendon lengths from all trials was calculated. Tendon length was determined as the two-dimensional distance between the AT insertion at the proximal portion of the calcaneus and the MTJ of GAM muscle (figure 4). To measure tendon length correctly, heel and MTJ displacements were transformed to a common coordinate system (CS). This was because the heel moved within the laboratory CS whereas the MTJ moved in the US probe CS. The principle of different coordinate systems is presented in figure 4. A similar method, although in three-dimensional space, has previously been used by Lichtwark and Wilson (2005). On this occasion two-dimensional analysis was considered adequate, because all movement could be restricted to a single plane. Kinematics and US videos were captured and digitized with Vicon Motus (v. 8.5, Vicon, Oxford, England) software. Tendon strain was calculated by dividing tendon elongation by tendon initial length.

Stiffness calculations and statistics

Tendon stiffness was deduced by using the least-squares method to find the line of best fit for ∆F vs. ∆L pairs from isometric contractions as shown in figure 5. Thus, the method takes into account the random variability in tendon force and length measurements because it is based on several pairs of data points. Young’s modulus was calculated by replacing tendon force with tendon stress and elongation with strain. To determine stiffness (and Young’s modulus) within the linear region, stress levels of 10MPa or above were accepted for analysis (Schechtman and Bader, 2002; Wang and Ker, 1995a). 10MPa stress usually corresponded to 10-20% of MVC level. The non-parametric Wilcoxon test was used to test for statistical differences and linear correlation between post/pre ratios of selected parameters was tested by calculating Pearson’s correlation coefficients (c.c.). The level of significance was set to P<0.05. In tables, mean, standard deviation (s.d.) and coefficient of variation (c.v.=s.d./mean) are reported. In figures, all values are means ± s.d.

Results

Subject anthropometric data and tendon mechanical properties before the exercise.

Tendon mechanical properties are displayed in table 1 and the main results of the study are summarized in table 2. Subjects performed 1150-2600 hops before exhaustion. Neither AT stiffness nor initial tendon length changed due to exercise. Tendon stiffness was 430±200 N mm–1 before and 390±190 N mm–1 after the exercise [not significant (n.s.)]. Despite the lack of changes in tendon stiffness, there was a significant reduction in maximum force production capacity from 3.5±0.6 kN to 2.8±0.7 kN (P<0.01).

Effect of exercise on mechanical properties of tendon.

A negative correlation was found between post-pre ratio of initial tendon length and post-pre ratio of stiffness (-0.622, P=0.037). Interpretation of this is that although neither stiffness nor initial length was significantly changed as a group, possible individual change in stiffness was related to initial length change, with the negative sign meaning that decrease in stiffness was observed together with tendon lengthening (creep). Stiffness also tended to decrease with increasing number of jumps but the correlation was non-significant (-0.527, P=0.118). There was no significant correlation between post-pre ratios of maximum tendon force and stiffness (0.423, P=0.223).

Figure 5 demonstrates a single ramp-derived stiffness against stiffness that was deduced from several ∆F vs. ∆L pairs over the linear strain region. Stiffness was independent of the method, being 117 N mm-1 (R=0.98, P<0.01) when derived from single ramp contraction and 118 N mm-1 (R=0.97, P<0.01) when derived from ∆F vs. ∆L pairs. This observation was repeated in all cases when ramp stiffness was calculated.

Individual Achilles tendon (AT) cross-sectional areas as a function of proximal distance from the calcaneus. Distance 0 cm corresponds to the first magnetic resonance image slice where AT is visible.

The individual tendon CSAs are plotted as a function of the proximal distance from the calcaneus in figure 6. Mean AT CSA was 43±11mm2 and lever arm length 4.9±0.4cm (table 1).

Discussion

The main finding of this study was that neither stiffness nor initial length of the GAM tendon changed after exhaustive high impact hopping exercise. Despite this, a significant reduction was observed in maximum isometric plantarflexion force (table 2). The results suggest a lack of acute mechanical changes in connective tissue regardless of diminished muscle force production capacity. The findings are in agreement with previous in vivo studies of human tendons that have not found an increase in tendon compliance after static or cyclic exercise (Mademli et al., 2006; Mademli et al., 2008; Ullrich et al., 2007) or after cyclic exercise alone ( Kubo et al., 2001b; 2005).

By contrast, some reports demonstrate that isometric contractions were able to increase VL tendon compliance (Kubo et al., 2001a; 2001b; 2005). This indicates that AT might be more fatigue resistant than VL tendon. Although some mechanical properties of human tendons are similar (Wren et al., 2001), fatigue resistance may be one property that varies. For example, high-stressed animal tendons showed higher fatigue resistance than low-stressed tendons despite equivalent Young’s modulus (Ker et al., 2000; Pike et al., 2000). The possibility that AT is more fatigue resistant than VL tendon has to be at least considered. Duration of loading could also be important. Total loading time was 500 s (10 s times 50 reps) in the isometric experiment by Kubo et al. (2005) whereas it was roughly the same in the current hopping task (0.25 s times 1885 reps). This suggests that longer duration isometric contractions could have more damaging potential. Interestingly, Wang et al. (1995b) showed the opposite; fatigue ruptures occur at shorter duration than would be predicted from creep ruptures alone. Similar results were reported by Noyes et al. (1974), who reported that connective tissue ruptures dominate over bone fractures at high strain rates. Nonetheless, natural movement rarely includes long duration loadings, and hence the current finding that AT was not fatigued after a single bout of hopping exercise is relevant with respect to human natural locomotion.

Wren et al. (2003) reported that in vitro tendons can be ruptured at physiological stress levels (40–60 MPa) after 1000 cycles or less if 8% initial strain level is exceeded. To estimate this, we assumed that maximal hopping stress roughly equals MVC stress. This was based on observations by Fukashiro et al. (1995), who deployed a buckle force transducer technique to demonstrate that peak tendon force (3.8 kN) is very similar to the current MVC force (3.5 kN) during low intensity hopping (7 cm) such as here (5–20cm). Therefore, calculation of maximum hopping strain yields an average of 5%. This is not enough to exceed the level set by Wren et al. (2003), but considering that strain in free tendon is higher (even double) than strain in tendon and aponeurosis combined (which was measured in the present study) (Magnusson et al. 2003; Finni et al. 2003), there is a good chance that the 8% limit was exceeded in the free tendon. Overall, calculations suggest that the stress and strain conditions set by Wren et al. (2003) to rupture AT with less than 1000 cycles were probably met in the present study.

Wren et al. (2003) are the only ones to demonstrate such short times to rupture. Also, they are the only ones to report induced rupture for human AT. Cycle times to rupture at comparable stress levels were much higher in tendons of sheep (Pike et al., 2000) and wallaby (Wang et al., 1995b), and also in human EDL tendons (Schechtman and Bader, 1997). The reason for different results could be methodological and due to difficulties in clamping techniques (Ker et al., 2000). Wren et al. (2003) note that the AT is especially hard to work with and easily leads to non-uniform strain. Furthermore, what makes comparison difficult is that there is not always a clear indication of tendon strain during fatigue tests (Pike et al., 2000; Wang et al., 1995b). Another disadvantage of cadaver tendons is that they are usually extracted from older donors and therefore suffer from possible tissue degradation and decreased stiffness due to ageing [for review of ageing, see Reeves (2006)].

Achilles tendon mechanical properties

Calculated AT force is prone to error in ankle moment arm length and assumptions of how force is distributed among plantar flexor tendons. It has been demonstrated that tendon moment arm changes due to ankle rotation and muscle contraction (Rugg et al., 1990; Maganaris et al., 2000) and that plantar flexion torque is not transmitted by AT only but has a contribution from deep plantar flexors (Finni et al., 2006, Finni et al. 2000). These possible sources of error are acknowledged here but no compensatory calculations are made, because not enough data is available to reliably estimate the amount of correction that should be used. In any case, a review of the forces in this study shows that they are in good agreement with directly measured tendon forces in vivo. Mean maximum isometric force (3.5kN) was higher than 1-3 kN in walking (Finni et al., 1998), approximately equal to 4kN in hopping (Fukashiro et al., 1995) and lower than 9kN in running at 16 km h-1 (Komi et al., 1992).

Tendon mechanical properties, such as stiffness or Young’s modulus, also rely on the ability to determine tendon elongation or strain. The average maximal tendon strain in this study (6.2%) is within the range given by others (Magnusson et al., 2003; Maganaris and Paul, 2002; Lichtwark and Wilson, 2005). Average tendon stiffness (430 N mm-1) matches previously reported values under similar conditions by Rosager et al. (2002), but is somewhat higher than reported during one-legged hopping (Lichtwark and Wilson, 2005) and lower than calculated for free tendon during isometric contractions (Magnusson et al., 2003). Maganaris and Paul (2002) used a different approach to estimate the contribution of gastrocnemius to the net torque and yielded a stiffness of 150 N mm-1. Previous reports on Young’s modulus have varied around 0.8-1.2 GPa (Magnusson et al., 2003; Rosager et al., 2002; Maganaris and Paul, 2002; Lichtwark and Wilson, 2005) and are lower than the current mean (2.0GPa). Differences could be induced by different methodology, because Lichtwark and Wilson (2005) used ultrasonography instead of MRI, and Magnusson et al. (2003) measured CSA at a constant distance (2 cm proximal) from calcaneus instead of using the smallest value. Both groups ended up with a higher tendon CSA, and this could explain their lower value for Young’s modulus.

Previous in vitro reports of Young’s modulus have determined a range 0.7-1.1GPa for human specimens (Schechtman and Bader, 1994; Johnson et al., 1994; Wren et al., 2001) but the comparison to in vivo is not straightforward. The strain value in this study includes free tendon and aponeurosis, and strain of free tendon has previously been shown to be much higher than for aponeurosis in active contractions (Finni et al., 2003; Magnusson et al., 2003). Therefore, strain in this study might be an underestimate (and Young’s modulus an overestimate) compared with in vitro.

Intra- and inter-individual variation

Figure 5 illustrates typical intra-individual variation, whereby some ∆F vs. ∆L pairs lie outside the predicted line. These random variations were treated by determining tendon stiffness as the slope of the best-fit line to several ∆F vs. ∆L pairs. To see how this method compares with stiffness derived from a single ramp contraction, some ramps were analysed point-to-point. An example of the comparison is shown in figure 5; pre-exercise stiffness that was determined from ∆F vs. ∆L pairs was 118N mm–1 (R=0.97, P<0.01) while stiffness determined from 80% ramp was 117N mm–1 (R=0.98, P<0.01). The former method was used because it allows analysis of more contractions and hence decreases the possible effect of random error. The methodological origin of trial-to-trial variation could be in the image tracking, because MTJ cannot always be unambiguously defined, especially after exercise when muscles sometimes become swollen and the visual appearance of the US image changes. Physiological variation could be caused, for example, by different activation of plantar flexor antagonists.

Tables 1 and 2 indicate that while the correlation values for each line fit to determine stiffness are significant, there is still high inter-individual variation in Young’s modulus (c.v.=0.50). This does not necessarily reflect such a high variety in tendon mechanical properties. More likely, it is due to the high variation in tendon maximal strain (c.v.=0.41). As mentioned earlier, free tendon has much higher strain, and therefore individuals with a higher proportion of free tendon tend to exhibit higher GAM tendon strains.

Conclusions

Even though the maximal force production capacity of young and physically active individuals was reduced (-20%) after a single bout of high impact hopping exercise (1150–2600 cycles), GAM tendon was not fatigued as evidenced by its unchanged stiffness. As a high-stressed tendon, AT might be more fatigue resistant than number of other tendons in the human body. Although results agree with many previous in vitro studies (Pike et al., 2000; Schechtman and Bader, 1997; Wang et al., 1995b,) they disagree with the one and only study conducted with human ATs (Wren et al, 2003). It is possible that clamping problems, often associated to in vitro experiments, could cause premature failure and thus different conclusions about tendon fatigue resistance between in vitro and in vivo. The current observation that tendon was not fatigued seems natural; it proves that the body is able to protect tendon from damage even during high impact exercise.

Ackowledgments

The authors gratefully acknowledge the technical staff at the University of Jyväskylä for assistance with this study. This study was supported by personal grant of the E. & A. Nyyssönen-foundation (J. P.). The authors have no conflict of interest to report.

List of abbreviations

AT,  Achilles tendon
CS,  coordinate system
CSA,  cross-sectional area
EDL,  extensor digitorum longus muscle
GAM,  gastrocnemius medialis muscle
MRI,  magnetic resonance imaging
MTJ,  myotendinous junction (most distal part of GAM muscle under US probe)
SOL,  soleus muscle
TA,  tibialis anterior muscle
US,  ultrasound / ultrasonography
VL,  vastus lateralis muscle
ΔF, tendon force
ΔL, tendon elongation

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