Diagnostics of Human Forearm Soft · i i i i i i i i JARKKO IIVARINEN Diagnostics of human forearm soft tissues using indentation and suction measurements Experimental and modeling

Publications of the University of Eastern Finland Dissertations in Forestry and Natural Sciences

Publications of the University of Eastern Finland

Dissertations in Forestry and Natural Sciences

isbn 978-952-61-1491-0

Jarkko Iivarinen

Diagnostics of Human Forearm Soft Tissues Using Indentation and SuctionMeasurementsExperimental and Modeling Analysis

Biomechanical methods for sensitive

analysis of mechanical properties

of soft tissues can be valuable in

diagnosing and monitoring the state

of pathologies, such as lymphoedema,

and aiding curative treatments. At

present, the relative contribution

of different soft tissues (skin, fat,

muscle) to mechanical response is

not well understood. This thesis

introduces new innovative modeling

methods to quantitatively analyze

soft tissues and measurement

instrumentation that may help to

diagnose and monitor alterations in

mechanical properties of soft tissues.

dissertatio

ns | 141 | Ja

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Iivar

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| Diagn

ostics of Hu

man F

orearm S

oft Tissu

es Usin

g Inden

tation and S

uction M

easurem

ents

Jarkko Iivarinen

Diagnostics of Human ForearmSoft Tissues Using Indentation

and Suction MeasurementsExperimental and Modeling Analysis

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JARKKO IIVARINEN

Diagnostics of humanforearm soft tissues usingindentation and suction

measurementsExperimental and modeling analysis

Publications of the University of Eastern FinlandDissertations in Forestry and Natural Sciences

No 141

Academic DissertationTo be presented by permission of the Faculty of Science and Forestry for publicexamination in the Auditorium L22 in Snellmania Building at the University of

Eastern Finland, Kuopio, on 27.06.2014,at 12:00 o’clock.

Department of Applied Physics

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Grano Oy

Kuopio, 2014

Editors: Prof. Pertti Pasanen, Prof. Kai Peiponen

Prof. Pekka Kilpeläinen, Prof. Matti Vornanen

Distribution:

University of Eastern Finland Library / Sales of publications

P.O. Box 107, FI-80101 Joensuu, Finland

tel. +358-50-3058396

http://www.uef.fi/kirjasto

ISBN: 978-952-61-1491-0 (printed)

ISSNL: 1798-5668

ISSN: 1798-5668

ISBN: 978-952-61-1492-7 (pdf)

ISSN: 1798-5676

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Author’s address: University of Eastern FinlandDepartment of Applied PhysicsP.O.Box 162770211 KuopioFINLANDemail: [email protected] / [email protected]

Supervisors: Dean Jukka S. Jurvelin, Ph.D.University of Eastern FinlandDepartment of Applied PhysicsKuopio, FINLANDemail: [email protected]

Associate Professor Rami K. Korhonen, Ph.D.University of Eastern FinlandDepartment of Applied PhysicsKuopio, FINLANDemail: [email protected]

Adjunct Professor Petro Julkunen, Ph.D.Kuopio University HospitalDepartment of Clinical NeurophysiologyKuopio, FINLANDemail: [email protected]

Reviewers: Associate Professor Cees Oomens (NL), Ph.D.Eindhoven University of TechnologyDepartment of Biomedical EngineeringP.O.Box 5135600 MB EindhovenTHE NETHERLANDSemail: [email protected]

Professor Jari Hyttinen (Finland), Ph.D.Tampere University of TechnologyELT / BioMediTechP.O.Box 69233101 TampereFINLANDemail: [email protected]

Opponent: Professor Arthur Mak (Hong Kong), Ph.D.Chinese University of Hong Kong, Shatin, N.T.Department of Biomedical EngineeringRm. 429, Ho Sin Hang Engineering BuildingHong Kong SARCHINAemail: [email protected]

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ABSTRACT

Several pathological conditions, such as soft tissue oedema, inducevariations in the mechanical properties of soft tissues. A devicewhich could sensitively detect these variations would be valuable indiagnosing and monitoring the pathology and aiding curative treat-ments. Hence, compact, hand-held devices, based on applicationof compression or external negative pressure (suction), have beenintroduced to detect changes in soft tissue stiffness. At present,little is known about the relative contribution of different soft tis-sue components (i.e., skin, adipose tissue and muscle) to the overallbiomechanical response. For the measurement device, impact ofthe shape and size of the contact head on the sensitivity of the mea-sured response is also usually unknown.

The present study evaluated the biomechanical role of differentsoft tissues in the relaxed, physically stressed (isometric flexor andextensor loading) and oedemic (venous occlusion) human forearm.Furthermore, the effects of changes in the thickness of the forearmtissues and device geometry on the sensitivity of the device wereanalyzed. The soft tissue response of the forearms in healthy hu-man subjects was measured under compression and negative pres-sure. At the site of mechanical loading, the geometry of differenttissue layers was measured using ultrasound (US) imaging and pe-ripheral quantitative computed tomography (pQCT). Finite element(FE) models were created and the model responses were matchedwith those obtained from the experimental measurements to de-termine the mechanical characteristics of the tissues in the model.Parametric sensitivity analyses were conducted to analyze the im-pact of tissue stiffness and thickness and the device geometry onthe model responses. Finally, two experimental negative pressuretreatment protocols (continuous and cyclic) were simulated to eval-uate their influence on the interstitial fluid flow in soft tissues.

The FE models could reproduce the experimental responses byincluding fibril structure of the skin, as well as the stiffness of theskin and adipose tissue. The simulated responses were more sensi-

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tive to changes in the mechanical properties of skin than to changesin adipose tissue and muscle. Skin thickness affected the sensitivityof the indentation instrument in detecting changes in the stiffnessof the underlying soft tissues.

During short-term oedema (venous occlusion), the reduction ofthe stretch of forearm tissues under negative pressure was relatedto stiffening of the skin and adipose tissue. In the simulation, it wasfound that continuous negative pressure therapy was more efficientin interstitial fluid transportation than the cyclic negative pressuretherapy.

In conclusion, the present measurement instrumentation canhelp to diagnose and monitor in vivo mechanical properties of softtissues in the forearm, especially those originating in the skin. Inthe measurements, it seemed that the skin predominantly controlledthe experimental responses. It seems that continuous negative pres-sure that transports the interstitial fluid effectively may help to re-duce pathological tissue swelling. Finite element modeling, espe-cially using fibril-reinforced models, proved to be a feasible methodto quantitatively optimize the geometrical aspects of measurementdevices. Further, these models could be used to study the long-termimpact of negative pressure therapies.

National Library of Medicine Classification: QT 34.5, WE 820, WR100, QS 532.5.A3, WE 500

Medical Subject Headings: Biomechanical Phenomena; Forearm; Skin;Adipose Tissue; Muscles; Collagen; Pressure; Suction; Isometric Contrac-tion; Edema; Extracellular Fluid; Finite Element Analysis; Sensitivity andSpecificity

Yleinen suomalainen asiasanasto: biomekaniikka; kyynärvarret; pehmytku-dokset; iho; rasvakudokset; lihakset; kollageenit; paine; imu; isometri-nen supistus; turvotus; solunulkoinen neste; elementtimenetelmä; sensiti-ivisyys

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Preface

This study was carried out during the years 2010-2014 in the De-partment of Applied Physics at the University of Eastern Finland.

First, I would like to thank all my supervisors for their guidanceduring this thesis project. Without your support and encourage-ment, this thesis might have never been finished. My principal su-pervisor Dean Jukka Jurvelin has been exceptional for his ability tothink out of the box and to find solutions for nearly every problem-atic dilemma that I faced during this project. My second supervisor,Associate Professor Rami Korhonen, has been especially encourag-ing and supporting with his hands-on approach to the challengesthat cropped up during my research. My third supervisor, AdjunctProfessor Petro Julkunen, has provided me with his never-endingsupport and enthusiasm in meeting the challenges encountered inmy research.

I am grateful to the official reviewers of this thesis, AssociateProfessor Cees W.J. Oomens, Ph.D. and Professor Jari Hyttinen,Ph.D., for their encouraging and constructive comments. I wouldalso like to thank Ewen MacDonald, D.Pharm., for linguistic review.

I thank all the brave researchers who volunteered to be my spec-imens in this study. I wish to thank Toni Rikkonen for support onpQCT imaging, Petri Sipola for help with the ultrasound imagingand Jari Arokoski for support with the muscle force meter and pres-sure cuff. I thank HLD Healthy Life Devices (Finland) for providingme the negative pressure treatment device used in this thesis. CSC -IT Center for Science (Finland) is acknowledged for computationalresources and Delfin Technologies Oy Ltd (Finland) for technicalsupport. Finally, I thank Ville-Pekka Vuorinen and Olavi Airaksi-nen for conducting the clinical pilot follow-up study.

I want to thank both Jukka’s BBC (Biophysics of bone and carti-lage) group and Rami’s CTB (Cell and tissue biomechanics) groupand my roommates Chibuzor Eneh and Xiaowei Ojanen for the

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companionship and interesting conversations. Mika Mononen is ac-knowledged for providing constructive criticism on the artwork ofmy publications and technical insights into the challenges in mod-eling and the use of LATEX and Janne Mäkelä because of reasons.Mika Mikkola is acknowledged for technical assistance on modelcreation. Pasi Miettinen and Antti Jaatinen are acknowledged fortechnical expertise and insight concerning life and matter. Joni-Pekka Pietikäinen is acknowledged for his assistance in interpret-ing other publications. I am grateful to Tarja Holopainen and othersecretaries of the Department of Applied Physics for supportingme through all the bureaucracy. In addition, I want to thank allmy friends who have supported me during these past years, es-pecially those in Internet Relay communication and Chat channels#fy2, #salakrillit and #hooseepaniikki. Furthermore, I wish to thankTBDP (National Doctoral Programme of Musculoskeletal Disordersand Biomaterials) for the annual meetings.

I wish to express my dearest thanks to my wife Henna. Also, Iwish to thank my daughter Sofia for all the grins and giggles. Fi-nally, I am grateful to my parents Jorma and Marja-Liisa, my step-father Tapani, my siblings Sanna and Kimmo and my twin brotherJuha for their support during the Ph.D. thesis project.

For providing financial support, the Finnish Funding Agencyfor Technology and Innovation (TEKES, funding decision 70011/09),North Savo Regional fund of Finnish Cultural Foundation, Instru-mentarium Science Foundation, Northern Savo Hospital District(EVO - special state subsidies), Doctoral Programme in MedicalPhysics and Engineering of University of Eastern Finland, Oy Lab-Vision Technologies Ltd (Lappeenranta, Finland), Foundation forAdvanced Technology of Eastern Finland and Kuopio UniversityFoundation are acknowledged.

Jarkko IivarinenKuopio, 02.06.2014

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To Henna and Sofia

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ABBREVIATIONS

2D Two-dimensional3D Three-dimensionalCLSM Confocal laser scanning microscopyFE Finite elementFEM Finite element methodFLVEL Fluid velocityFR Forearm at restFRHE Fibril-reinforced hyperelasticFRPHE Fibril-reinforced porohyperelasticHE HyperelasticIEL Isometric extensor loadingIFL Isometric flexor loadingLE Logarithmic strainMFM Muscle force meterMRI Magnetic resonance imagingMSE Mean squared errorMVC Maximal voluntary contractionnMSE Normalized mean squared errorPHE PorohyperelasticPOR Pore pressurepQCT Peripheral quantitative computed tomographys SubjectSA Symmetry axisSD Standard deviationSEM Scanning electron microscopeSLS Standard linear solidUS UltrasoundUV UltravioletVHE ViscohyperelasticVO Venous occlusion

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SYMBOLS

B0 Initial bulk modulusc Viscous damping coefficientC10 Shear parameter of neo-Hookean hyperelastic modelD1 Bulk parameter of neo-Hookean hyperelastic modele Void ratioE Elastic modulusE f Fibril network modulusF Forceg Acceleration due to gravitygi Relative shear relaxation modulusG Shear modulusG0 Initial shear modulusG∞ Long-term shear modulusI Unit tensorI1 First deviatoric strain invariant of neo-Hookean hyperelastic modelJel Elastic volume ratiok PermeabilitykH Spring constantki Isotropic permeabilityks Fully saturated permeabilityK Hydraulic conductivityKs Dependence of permeability on saturationn Porosityp Fluid pressureS Von Mises stresst TimeU Strain energy potential functionvw Fluid velocityVp Volume of poresVs Volume of solidsx Deformation

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α Nonlinearity parameter of Ogden hyperelastic modelβ Velocity coefficientγw Specific weight of the wetting liquidε Strainε0 Fibril toe limit strainη Newtonian viscosity coefficientηG Shear viscosity coefficientν Poisson’s ratioνk Kinematic viscosityρ Densityσ StressσE Effective solid stress tensorσ f Fluid stress tensorσfibr Stress tensor of fibresσnonfibr Stress tensor of nonfibrillar matrixσs Solid stress tensorσt Total stressτ Relaxation time

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LIST OF PUBLICATIONS

This thesis consists of the present review of the author’s work inthe field of computational modeling of human forearm soft tissuesand the following selection of the author’s publications:

I Iivarinen J.T., Korhonen R.K., Julkunen P. and Jurvelin J.S.,“Experimental and computational analysis of soft tissue stiff-ness in forearm using a manual indentation device”. MedicalEngineering & Physics. 33, 1245–53 (2011).

II Iivarinen J.T., Korhonen R.K., Julkunen P. and Jurvelin J.S.,“Experimental and computational analysis of soft tissue me-chanical response under negative pressure in forearm”. SkinResearch and Technology. 19, e356–65 (2013).

III Iivarinen J.T., Korhonen R.K. and Jurvelin J.S., “Modeling ofinterstitial fluid movement in soft tissue under negative pres-sure – relevance to treatment of tissue swelling”. (SubmittedApr 2014).

IV Iivarinen J.T., Korhonen R.K. and Jurvelin J.S., ‘Experimentaland numerical analysis of soft tissue stiffness measurementusing manual indentation device – significance of indentationgeometry and soft tissue thickness”. Skin Research and Technol-ogy. 1-8 (Epub 2013 Nov 23).

Throughout the overview, these papers will be referred to by Ro-man numerals.

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AUTHOR’S CONTRIBUTION

The publications in this dissertation are original research papers onexperimental and simulation studies of human soft tissues. The au-thor has been the main contributor to each method presented inthe publications and carried out all of the data analyses and sim-ulations. The author conducted all the mechanical and imagingmeasurements. The author has been the main writer of each paper.

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Contents

1 INTRODUCTION 1

2 SOFT TISSUES 32.1 Skin . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 32.2 Adipose tissue . . . . . . . . . . . . . . . . . . . . . . . 72.3 Muscle . . . . . . . . . . . . . . . . . . . . . . . . . . . 9

3 BIOMECHANICAL PROPERTIES OF SOFT TISSUES 133.1 Skin . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 133.2 Adipose tissue . . . . . . . . . . . . . . . . . . . . . . . 173.3 Muscle . . . . . . . . . . . . . . . . . . . . . . . . . . . 183.4 Changes in soft tissue characteristics . . . . . . . . . . 19

4 CONSTITUTIVE MODELING OF SOFT TISSUES 214.1 Elasticity . . . . . . . . . . . . . . . . . . . . . . . . . . 224.2 Viscoelasticity . . . . . . . . . . . . . . . . . . . . . . . 234.3 Poroelasticity . . . . . . . . . . . . . . . . . . . . . . . 254.4 Fibril-reinforcement . . . . . . . . . . . . . . . . . . . . 26

5 AIMS OF THE THESIS 29

6 MATERIALS AND METHODS 316.1 Subjects . . . . . . . . . . . . . . . . . . . . . . . . . . . 316.2 Ultrasound measurements . . . . . . . . . . . . . . . . 326.3 pQCT and segmentation . . . . . . . . . . . . . . . . . 326.4 Mechanical measurements . . . . . . . . . . . . . . . . 33

6.4.1 Compression . . . . . . . . . . . . . . . . . . . 356.4.2 Negative pressure . . . . . . . . . . . . . . . . 36

6.5 FE models for skin, adipose tissue and muscle . . . . 376.5.1 Model creation and geometry . . . . . . . . . . 376.5.2 Elements and boundary conditions . . . . . . 396.5.3 Loading protocol . . . . . . . . . . . . . . . . . 40

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6.5.4 Material model description . . . . . . . . . . . 416.5.5 Sensitivity analysis and optimization . . . . . 42

7 RESULTS 457.1 Indentation . . . . . . . . . . . . . . . . . . . . . . . . . 45

7.1.1 Experimental measurements . . . . . . . . . . 457.1.2 Simulation . . . . . . . . . . . . . . . . . . . . . 45

7.2 Negative pressure . . . . . . . . . . . . . . . . . . . . . 487.2.1 Experimental measurements . . . . . . . . . . 487.2.2 Simulation . . . . . . . . . . . . . . . . . . . . . 48

8 DISCUSSION 558.1 Experimental measurements and soft tissue diagnostics 558.2 Model optimization . . . . . . . . . . . . . . . . . . . . 568.3 Sensitivity analysis . . . . . . . . . . . . . . . . . . . . 57

8.3.1 Tissue stiffness . . . . . . . . . . . . . . . . . . 578.3.2 Tissue thickness . . . . . . . . . . . . . . . . . . 598.3.3 Stiffness meter geometry . . . . . . . . . . . . 60

8.4 Fluid flow . . . . . . . . . . . . . . . . . . . . . . . . . 608.5 Validity and limitations of models . . . . . . . . . . . 628.6 Clinical study - preliminary results . . . . . . . . . . . 65

9 SUMMARY AND CONCLUSIONS 67

REFERENCES 69

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1 Introduction

The stiffness of soft tissues is known not only to vary from siteto site but also to change with age as well as due to pathologiesand malignancies. A device with the capacity to sensitively moni-tor changes in soft tissue stiffness would be valuable in diagnosingand monitoring the state of pathologies and perhaps could help inthe search for curative treatments. In fact, several hand-held in-strumentation devices have been claimed to detect changes in softtissues. However, very little is known about the relative contribu-tion of each soft tissue component (i.e., skin, adipose tissue andmuscle) to the overall biomechanical response. Furthermore, howthe shape and size of instrumentation influences the sensitivity ofthe measured response is usually unclear. A realistic model for thehuman soft tissues could help to optimize the design of an experi-mental instrumentation as well as improving the protocols for thetreatment of pathologies, e.g., oedema and lymphoedema.

Human soft tissues are complex materials, often presenting fibre-reinforced, nonlinear, viscoelastic and anisotropic mechanical prop-erties [1–15]. Several techniques have been used to investigate softtissue mechanics, e.g., compression [1, 3, 10, 16–24] and negativepressure [3,5,25–28] measurements. The elastic modulus or varioushyperelastic model parameters for the tissues are often reported.The finite element method (FEM) has been used to study the me-chanical characteristics and parameters of tissues [13, 15, 27, 29–35],the biomechanics of tactile sensation [36], the stress distributionin gluteal [4, 10, 37–39], plantar [11, 40, 41] and subcutaneous adi-pose [12] tissue, skin wrinkling [42], the relative contributions ofskin layers to the experimental response [43], collagen fibre dis-persion and contribution [44], breast deformation under gravita-tional force and compression [45–47], deep tissue injury in a trans-tibial amputee [19], muscle contraction [48–50], muscle under im-pact loading [51], intramuscular pressure [52] and muscular hy-

Dissertations in Forestry and Natural Sciences No 141 1

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Jarkko Iivarinen: Mechanical diagnostics of human forearm soft tissues

drostats [53]. There are no 3D models of human forearm whichincorporate features such as fibril-reinforced hyperelastic skin, andhyperelastic adipose tissue and muscle; these features are not takeninto account when evaluating tissue deformation under compres-sion and negative pressure. In addition, there are no earlier modelsof the human forearm which have investigated the effect of negativepressure treatment on fluid flow. Even though finite element (FE)modeling has frequently been used in the design of bioengineeredmaterials, e.g., aortic valves, coronary stents, hip stems, meniscalimplants and voice-producing prostheses, much less optimizationof the measurement devices has been conducted [54].

The aim of the present study was to investigate the sensitiv-ity of the studied instrumentation towards changes in the tissuesand in the geometry of the device. A further aim was to evalu-ate the impact of physical stress and short-term oedema in the softtissues and to compare different negative pressure treatment pro-tocols. The biomechanical role of different soft tissues in relaxed,physically stressed and oedemic (venous occlusion) human forearmwas evaluated. The impact of changes in the thickness of tissuesand device geometry on the sensitivity of the device was analyzed.Layered, hyperelastic FE models were created and the model re-sponses were matched with those obtained from the experimentalmeasurements in order to determine the mechanical characteristicsof the tissues in the model. The mechanical roles of the tissues andthe performance of the instruments were evaluated by conductingparametric sensitivity analyses taking into account the impact ofthe changes in the stiffness and thickness of the tissues and in thedevice geometry on the model responses. Finally, two experimentalnegative pressure treatment protocols (continuous and cyclic) wereexamined to evaluate their impact on interstitial fluid flow in softtissues.

2 Dissertations in Forestry and Natural Sciences No 141

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2 Soft tissues

The musculoskeletal system provides human body with form, sup-port and the ability to move [55]. Musculoskeletal system is entirelycovered by layers of soft tissues [1,56]. These soft tissue layers typi-cally consist of skin and adipose tissue (Figure 2.1), which has phys-iological importance, e.g., protecting the body from mechanical andchemical hazards and regulating temperature and water loss. Thesize, thickness and properties of soft tissues may vary due to age,gender, human race, ultraviolet (UV) radiation, physical exerciseand nutrition.

2.1 SKIN

The human skin has a complex structure and it is the heaviest andthe largest organ in the human body [3] with an area of 1.5 - 2 m2

and a weight of 5 - 25 % of the total body weight [57, 58]. It isthe interface between the external environment and the body and,consequently, has a number of important functions [59], e.g., pro-tecting the organs from potentially injurous hazards (mechanicalstresses, electromagnetic radiation, harmful organisms and chemi-cals) [58,60], but also temperature regulation (perspiration and con-trol of blood flow in capillaries) and vitamin D synthesis [61, 62].Furthermore, the skin possesses various nerve endings and sensoryreceptors [60] to help in the detection of pain and temperature andalso specialized glands for the generation of sweat and sebum [59].Skin is also the site where interstitial fluid and lymph are formed.Human skin is a multilayered material [42, 63] with well-definedanatomical regions [64]. Each layer has a distinctive structure andcharacteristic properties [60, 65]. According to the classical viewon the skin, the layers are epidermis, dermis and hypodermis (adi-pose tissue). However, nowadays the skin is usually considered astwo-layered [11, 60] and the adipose tissue is treated as a separate


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Figure 2.1: Schematic representation of the soft tissues of forearm. The skin consists of (A1)epidermis and (A2) dermis, (B) adipose tissue, (C) muscle, (D) artery, (E) lymphatic vessel,(F) vein, (G) nerve and (H) fascia. Image modified and reproduced with kind permissionof Skin Care Forum, BASF Personal Care and Nutrition GmbH, http://www.skin-care-forum.basf.com.

organ [66–68]. This classification is also followed in the present the-sis. The normal thickness of skin is around 1 - 4 mm (excluding theadipose tissue) [58].

The epidermis is an epithelial layer [60] with thickness of 30- 640 µm [69–71]. The epidermis has a surface layer composedof dead cell tissue (stratum corneum) [59, 72] and a viable epider-mis layer underneath. The stratum corneum is around 10 - 40 µmthick [59, 65, 69] and contributes little to the mechanical propertiesof the skin [64]. These epithelial cell layers are almost impermeable


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Soft tissues

to water and are chemically inert [60], and therefore they preventwater loss and protect the body from external contaminants [65].Underneath the stratum corneum one finds the viable epidermis,which mainly consists of keratinocytes [72], specialized cells that re-new the skin continually. Therefore, epidermis has a high capacityfor regeneration after damage [60] as the cells of viable epidermisreplace those of the dead cells of the stratum corneum as they areworn away [65]. The vascular supply of the skin is limited entirelyto the dermis and therefore the epidermis relies on the capillaries inthe dermis for its supply of nutrients and metabolic exchange [60].The junction of the epidermis with the underlying dermis adheresin a undulating manner, such that large cones of epidermal tissuepenetrate through the dermis [72] (Figure 2.1).

The dermis (corium) is a 0.5 - 4 mm thick layer beneath theepidermis [58, 64, 70, 73]. The dermis generates the appendaces ofthe skin (hairs, nails, sudorific and sebaceous glands) [74]. Fur-thermore, nerve endings (mechanoreceptors) and blood and lymphvessels of the skin are located in the dermis [75]. The primaryfunctions of the dermis are to provide nourishment and mechan-ical support to the epidermis and to store water [73, 76]. Dermis iscomposed of papillary and reticular layers, which differ mostly inthe properties of collagen [11, 64]. The papillary dermis comprisesabout 10 % of the full dermis thickness [64]. Dermis is mostly com-posed of fibrous proteins collagen and elastin in the interfibrillarmatrix [2, 59, 65, 77, 78]. The main component of the dermis is thecollagen (27 - 39 % of the volume, 75 - 80 % of the dry weight) [5,73].The remaining components of the dermis consist of elastin (0.2 - 0.6% of the volume, 4 % of the dry weight), glycosaminoglycans (0.03- 0.35 % of the volume) and water (60 - 72 % of the volume) [73].

Collagen is a basic structural element for soft and hard tis-sues [79] and is present in every connective tissue, where it pro-vides mechanical integrity and strength. Collagens are the mostabundant family of proteins in the human body and account forapproximately 25 % of the total dry weight in mammals and asmuch as 77 % of the dermis [80, 81]. The amino acids in collagen


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are mostly in a triple-helical conformation [79]. Dermis containsprimarily type I and III collagen fibrils of 20 - 100 nm in diame-ter that are packed into thicker collagen fibres of 0.3 - 40 µm indiameter [60, 64, 73, 77, 80, 82]. Type III collagen accounts for 15 %of dermal collagen [64]. The collagen fibres in dermis are heavilycrimped (wavy, undulating) and are not oriented in parallel as theyare in tendons or cartilage [65]. However, the general orientationof the collagen fibres in the dermis depends on the location withinthe body and is described by cleavage (“Langer”) lines [11, 83, 84].These lines correspond to those of maximum tension [85]. Elastinis a rubberlike fibrous protein that forms a meshlike network indermis [73] where it provides elasticity for soft tissues and smooth-ness for skin [79]. Elastin microfibrils are 10 - 12 nm in diameterand they form an interwoven rope-like structure of elastin fibreswith diameters of 1 - 4 µm [64, 73, 77]. The interfibrillar matrix,“ground substance” surrounding the fibrous components of der-mis, is composed of glycosaminoglycans (mostly dermatan sulfate,hyaluronic acid and chondroitin sulfate), proteoglycans and glyco-proteins [73, 77].

The walls of the vascular system permit free diffusion of fluidand small organic molecules [60]. Interstitial fluid is formed whenblood is seeped through capillaries into the interstitium. Most ofthe interstitial fluid passes back into the blood capillaries and someof the interstitial fluid passes to the lymphatic system. The lym-phatic system consists of lymphatic vessels and lymphatic nodes[86]. Lymph fluid is formed when the interstitial fluid flows intoneighbouring initial lymphatic capillaries [60, 86]. Blind-endinglymphatic capillaries terminate in the dermis near to the base ofepidermis and drain into a lymphatic vessel network at the junc-tion of the dermis and adipose tissue [60, 86, 87]. Lymph can con-tain plasma proteins, immune cells, infectious organisms and anti-gens [88, 89]. The lymphatic vascular system is essential in the reg-ulation of tissue volume and pressure (homeostasis) by transport-ing excess fluid from the interstitial space to the blood circulation,and it also can be considered to aid the immune system [88–90].


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Soft tissues

The larger lymphatic vessels contain valves that prevent backwardflow [86, 88, 89]. All lymphatic vessels terminate in two large lym-phatic vessels (the thoracic duct and the right lymphatic duct) andreturn the lymph to the blood circulation [60, 86]. If lymph ves-sels become obstructed, the tissues drained by these vessels maybecome oedemic [60]. However, visible swelling (lymphoedema) isdetected only after the lymph flow has been reduced by 80 % [91].

2.2 ADIPOSE TISSUE

Adipose tissue can be found of varying locations of the humanbody, e.g., visceral adipose tissue (surrounding organs) and sub-cutaneous adipose tissue (often surrounding muscles) [68, 92–94].Adipose tissue may have special biological functions, such as mam-mary gland, heel fat pad and bone marrow adipose tissues [67].Adipose tissue concentrates much of the body’s fat around the ab-domen and breasts [12]. The structure and mechanical propertiesof adipose tissue vary greatly, especially in regions of attrition, suchas in the plantar, palmar, and digital areas [95–97]. Relatively smalldifferences in the adipose tissue composition exist at other bodysites [98]. The main adipose tissue compartment is the subcuta-neous adipose tissue [66]. In this study, the subcutaneous adiposetissue will be simply termed as adipose tissue.

Subcutaneous adipose tissue (subcutis, hypodermis) is a layerof loose connective tissue underneath the dermis [60, 75]. The dis-tibution of adipose tissue is influenced by gender, age, genotype,diet, physical activity level, hormones, and drugs [68]. In humans,adipose tissue appears in two forms: white (in adults) and brown(in infants) [99]. White adipose tissue functions as a mechanicalimpact absorber, load distributor, energy storage, thermal insula-tor [60,67,75,95] and as an endochrine - a source of hormones (lep-tin) [100–102]. The main function of brown adipose tissue is heatproduction (thermogenesis) [67, 68, 92]. The dominant componentof white adipose tissue is lipidic fluid (60 - 85 % of the weight,60 - 70 % of the volume) composing mostly of triglycerides, free


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fatty acids, diglycerides and cholesterol phospholipids [59, 98, 99].The remaining components are water (5 - 30 % of the weight) andprotein (2 - 3 % of the weight) [96, 98, 99]. Adipose tissue mass re-flects the net balance between the energy expenditure and energyintake [66, 92]. The adipose tissue mass in slim adult humans is 9 -18 % of body weight in men and 14 - 28 % in women [66, 68]. Theprimary phenotypic characteristic of obesity is the unhealthy ex-cess amount of white adipose tissue [68] which can increase up to4-fold in massively obese persons, reaching 60 - 70 % of total bodyweight [66].

Adipose tissue consists predominantly of adipocytes (lipocytes,fat cells) [12, 96, 103]. The adipose tissue is well vascularized. Eachadipocyte is in contact with at least one capillary [92, 98, 104]. Thewhite adipocytes are spherical and filled with a large fat droplet[96,98]. The fat droplet functions as an energy deposit [95]. The di-ameter of the white adipocytes is 30 - 70 µm [92]. In healthy adults,one third of the adipose tissue is comprised of mature adipocytesand the remaining two thirds of blood vessels, nerves and fibrob-lasts [92]. When there is weight gain, there is an initial increasein adipocyte volume until a “critical” tissue mass is reached, af-ter this point recruitment of new adipose cells takes place [105].In weight loss, there is a reduction in both adipocyte number andvolume [106].

The subcutaneous adipose tissue is formed by two layers: onemore superficial in contact with the dermis (areolar layer) and theother directly beneath (lamellar layer) [92,94,95,107–109]. Lamellarlayer is unique in its potential to undergo an enormous volumechange [66] which increases in thickness when an individual gainsweight [95,107,108]. The areolar layer does not change in thicknessas the lamellar [95]. The lamellar layer includes the lymphatic andblood vessels [95]. In obese patients, the lamellar layer may be8 - 10 times thicker than that in normal-weight people, while theareolar layer may only double in thickness [95]. The areolar andlamellar layers are separated [95, 107, 108] and the skin is anchoredto the underlying bones or the deep fascia of muscle by superficial


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Soft tissues

fascia [73, 110]. Superficial fascia is a loosely interwoven weblikecollagen fibre network in the adipose tissue, where adipocytes aregathered [60, 96, 98].

2.3 MUSCLE

Muscle accounts for about 40 - 50 % of an adult human’s totalweight [111]. It ensures the ability to move, and it provides me-chanical support and can generate heat [112–114]. Nutrition andphysical exercise influence the muscles. In addition, differences inmuscles occur between genders [115] and aging is associated withmuscle weakness [116]. Muscle is a composite tissue of contractilematerial, connective tissue, blood vessels and nerves [117]. Not onlythe movement of the body but also the flow of fluid in the vascu-lar system are generated by the contraction of muscles [52]. Thereare three muscle types: skeletal, heart (cardiac) and smooth [79].Only the contraction of the skeletal muscle is voluntary. Skeletaland heart muscles but not smooth muscle are striated (transver-sally striped) [79]. Skeletal muscle cell is multinucleated, the heartmuscle cell has 1 - 2 nuclei and smooth muscle cell has one nu-cleus [60].

Skeletal muscles are usually connected through tendons to thebones [52] and generate a force, leading to motion and they provideprotection to the underlying skeleton [50, 55, 79, 118, 119]. Skeletalmuscle represents the majority of the muscle tissue in the body[112]. Humans have around 700 different skeletal muscles [120].Skeletal muscle is capable to powerful contractions [112]. During acontraction, the cross-sectional area of the skeletal muscle increasesby 20 - 32 % [121] and simultaneously shortens maintaining a con-stant volume [113, 122]. Skeletal muscle is well-vascularized [73].The heart muscle has more mitochondria and capillary blood ves-sels than the skeletal muscle [79]. Skeletal muscle consists 75 - 80% of water, 10 - 20 % of proteins (collagenous tissues) and rest saltsand fat [14, 21, 123]. The diameter of an elongated skeletal musclefibre is 10 - 60 µm and length from less than 1 cm up to 30 cm [79].


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Heart muscle can only be found in heart and in the large vesselsclose to heart [112]. It is responsible for pumping blood throughoutthe blood vessels to the body. Heart muscle consists of branchingnetwork of interconnecting fibres (cells) [79] linked electrically andmechanically to act as a whole and work as a unit [79, 112]. There-fore, graded contraction does not occur in heart muscle, instead itis all-or-none [79]. Contractions of heart muscle are less power-ful than those of skeletal muscle but is resistant to fatigue [112].Smooth muscle can be found in the blood vessels (vascular), in theintestine (intestinal), associated with hair follicles in the skin and lo-cated in the eyeball [79, 112]. Smooth muscle consists of elongatedor spindle-shaped fibres [112]. The contraction of smooth muscle isslow, smooth and sustained [112].

Skeletal muscle is composed of multiple layers of nested bun-dles, each held together and shrouded via extracellular matrix pro-teins and connective tissue sheath [52,79], which fills the spaces be-tween the muscle fibres within a bundle [79]. The outermost con-nective tissue layer, wrapping whole skeletal muscle, is the deepfascia which also serves for attachment of the skin via the super-ficial fascia [58, 73]. Usually, the underlying muscles are free toglide beneath the deep fascia [73]. Skeletal muscle is composed offascicles containing parallel bundles of long fibres [52, 112]. Fas-cicles consist of muscle fibres which are composed of myofibrils[21, 55, 79]. Myofibrils are striated when they are stained by dyesand when they are examined optically [79]. The stripes in myofib-rils are the sarcomeres which are arranged in series [55], the small-est functional units of muscle [79]. Each myofibril is composed ofarrays of myofilaments [79]. Two types of overlapping myofilamentexist in sarcomere: actin and myosin molecules [55, 79, 118].

Skeletal and heart muscle contraction results from cyclic inter-actions between actin and myosin in which the chemical energy isconverted into mechanical work, force, and shortening [114]. Pas-sive force production occurs when a muscle has been stretched fromits resting length. The force is largely attributable to the extracel-lular matrix [55]. When the muscle tissue is stretched until myosin


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Soft tissues

and actin filaments no longer overlap, then the filaments generateno further tension [124]. Active force production consumes energy;this occurs when skeletal muscle activation causes overlapping ofthe actin and myosin filaments which interact via cross-bridging,resulting in muscle contraction [55]. During skeletal muscle con-traction, all of the muscle fibres are not excited at the same time [79].The stress induced by active skeletal muscle in the direction of thefibres is often described as the sum of an active and a passive com-ponent [21, 48, 125]. Therefore, the total force of contraction of askeletal muscle depends on how many muscle fibres are stimu-lated [79].


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3 Biomechanical properties ofsoft tissues

The fundamental aim of biomechanics is to understand and charac-terise the mechanical behaviour of biological tissues. The mechan-ical behaviour of biological materials is determined by their struc-ture and composition: macroscopic and microscopic tissue mor-phology, fluid flow in the tissue, conformation of fibres and the in-terfaces between various structures [1]. The biomechanical proper-ties of tissues are known to vary from site to site and to change withage and pathological conditions [126]. Furthermore, differences inthe properties of tissues exist in in vitro and in vivo analysis [127].It is difficult to accurately describe the mechanical behaviour of bi-ological tissues under loading [1,128]. Several testing methods andprotocols have been established as ways to characterize the tissuebehaviour but they often provide only problem-specific materialparameters [51]. Typical testing protocols involve determination ofmechanical properties in tension and compression [1]. Nowadays,mechanical measurements are often coupled with FE modeling toenhance the analyses on tissue characteristics.

3.1 SKIN

As human skin is a complex composite material consisting of fluidin a porous medium, fibrous collagen and elastin in ground sub-stance with viscoelastic (time-dependent) characteristics, highly re-alistic description of the mechanical behaviour of the skin is alsocomplex. Skin possesses fibre-reinforced, nonlinear, viscoelastic,nearly incompressible, anisotropic and heterogeneous mechanicalproperties [3, 5, 7, 8, 27, 32, 63–65, 76, 83, 84, 126, 129–132]. Each layerof the skin has different mechanical properties [42, 63, 132]. Inaddition, the skin is in tension in the natural “relaxed” state in


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vivo [25, 133]. This makes accurate predictions of the mechanicalbehaviour of the skin challenging [65].

The mechanical properties of human skin depend mainly on thecollagen and elastin fibre network and the ground substance of thedermis [26, 31, 64, 65, 76, 133, 134], with some contribution by theepidermis and the stratum corneum [26]. Collagen is the principalfibrous load bearing component in skin and provides considerablemechanical strength at high load levels [2,60,73,74]. Collagen in thedermis permits large deformations requiring low forces. However,it also exhibits a high tensile strength to limit tissue deformation toprevent rupture of the epidermis [76, 135]. The stratum corneumof epidermis can sustain, depending on its relative humidity, up to22 - 190 % strain before fracturing [136]. The anisotropic behaviourof the skin is mostly determined by the organisation of collagenfibres [5, 44, 132, 137]. It is important to be aware of the collagendistribution if one wishes to accurately describe the behaviour ofthe skin [44, 129], however, the concentration and orientation dataof the collagen is not often available. With respect to the knownmaterials, elastin best resembles an ideal linear elastic solid material[79] and it provides skin with its mechanical integrity at low loads[73]. Elastin can be considered as a return spring which contractsthe skin after the mechanical load is removed [135]. The mechanicalcontribution of the ground substance is small [129]. However, theground substance and water in porous medium contribute to theviscoelastic nature of skin [5, 73] in addition to the collagen [138,139].

The highly nonlinear mechanical response of the skin undertension can be explained by collagen fibre straightening [65]. Thenonlinear stress-strain response can be divided into three phases[64, 129, 135] (Figure 3.1). In the relaxed state, the collagen fibreshave crimped appearance [78]. In the first phase, the crimped col-lagen fibres starts to straighten, reorientating towards the load axisand contribute very little to the mechanical response [5, 65, 78, 129,135, 140], at this time the response is dominated by elastin fibresand ground substance [64, 73, 141, 142]. Variations in the degree


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Biomechanical properties of soft tissues

Figure 3.1: Three phases of stress-strain response and the orientation of the collagen of theskin under tension. Images of collagen taken with a scanning electron microscope (SEM).At the beginning of phase one skin is unstressed and the collagen is crimped. At the endof phase one, fibres have started reorienting towards tensional load axis while some fibresare already almost straight. At the end of phase two, the fibres are mostly aligned andnearly straight. During the phase three, fibres are fully aligned and straight. SEM imagesreproduced and modified from [135] with kind permission of John Wiley and Sons.

of collagen undulation exists, so the fibres begin to bear the loadduring increasing tension [129]. The soft tissues may roughly becategorized into two main groups: short and long phase one [143].In the second phase (often termed the toe region), collagen fibresprogressively straighten, align parallel and start to bear load, char-acterized by increasing tissue stiffness [64, 65, 129, 135, 139]. In thethird phase the collagen is mostly aligned and contribute highly tothe mechanical response of the skin [5, 129]. The response of theskin is almost linear, very little extension is possible and the stiff-ness increases rapidly [2, 64, 65, 129, 135, 140]. Deformation beyondthe third phase leads to collagen fibril defibrillation and rupture ofthe skin [64, 129, 135].

The biomechanical characteristics of skin have been extensively


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investigated. Skin properties have been measured under compres-sion [1, 3, 6, 7, 16, 22, 23, 32, 34, 35, 38, 42, 56, 72, 74, 126, 144–150], ten-sion [8, 25, 30, 54, 65, 76, 129, 132, 151–156], negative pressure load-ing (suction) [3, 5, 25–28, 31, 43, 63, 157–159], torsion [134, 160, 161],friction [59, 149, 162–166] and motion analysis [65, 167]. Perhaps,the simplest mechanical measurement method is the in vitro ten-sile testing, where the material constants can be determined un-der a carefully controlled enviroment [54]. Tissue indentation isalso a common technique used to investigate large deformations invivo [13]. When the skin surface is under mechanical loading in vivo,it is important also to consider the contribution by the deeper tissuelayers [63]. During a high suction stress, the effects of skin cannotbe isolated from those of the adipose tissue [63]. The resistance tothe applied vertical stress in vivo is essentially due to the skin ratherthan the adipose tissue, although the relative contribution of eachcannot be easily evaluated [63].

Skin tissue has been modeled as an elastic [54, 146] or a hyper-elastic material [2, 10, 19, 27, 31, 34, 45, 47, 65, 152, 168] but occasion-ally the material description has been extended with inclusion ofviscoelastic [1, 8, 42, 43, 169, 170], poroelastic [32], fibre-reinforced[11,41,44,78,131,171] or anisotropic [28] properties. Often, the elas-tic modulus or various hyperelastic and viscoelastic model param-eters are reported for the skin. However, skin has no distinctiveelastic modulus mostly due to collagen and the high dependenceon the applied strain [6, 79, 159, 172]. Consequently, values of elas-tic modulus of the skin can roughly be categorized into two maingroups: less than 100 kPa (range 0.3 - 100 kPa) [3, 23, 27, 56, 74, 76,144, 146, 149, 150, 153, 155, 161, 173, 174] and above 100 kPa (up to850 kPa) [5,8,26,30,134,157–159], where elastic modulus values areusually low for indentation and high for suction. The modulus,specifically of phase one, varies extensively in different reports, e.g.,from 5 kPa [76], 11.2 kPa (range 0.22 - 58.4 kPa) [56] up to 41 -131 kPa [5]. The increase in the modulus can be 5-fold as the skinprogresses from phase one to phase three [5]. The elastic modu-lus of the epidermis is relatively high, 0.49 - 2.6 MPa [72, 134, 145].


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However, as it is so thin, it may not contribute greatly to the overallstiffness of the skin [65]. Muscle contraction can induce a 4-fold in-crease in stiffness of the skin [16, 23, 150]. The fibril toe limit strain(deformation level at phase two when most collagen is aligned andstart to bear load) for skin in vivo during negative pressure may be3.8 - 6.8 % [5]. The elastic modulus of collagen has been reported tovary from 48 MPa up to 10 GPa [5,60,64,129,175,176]. The modulusis often low for the collagen network but high for a single collagenfibre.

3.2 ADIPOSE TISSUE

Subcutaneous adipose tissue is a widely neglected topic in biome-chanics and often ignored when the in vivo characteristics of humansoft tissues are being studied and simulated [12, 13, 93, 98, 99]. Usu-ally, investigation of the adipose tissue has focused on the heel fatpad and the breast tissue [12,98]. Adipose tissue is important in theload transfer between different structures in the body [12,96,98,177]and significantly influences the deformation of soft tissue compos-ite as a whole [12]. Adipose tissue exhibits nonlinear and viscoelas-tic properties [9, 18, 93, 96, 170] but it has been reported to displaylinear characteristics up to 50 % of tensile strain [93]. With respectto shear, adipose tissue is highly nonlinear for strains larger than0.1 % [98].

The mechanical properties of adipose tissue have been mea-sured under compression [1, 9, 11–13, 18, 20, 22, 33, 37, 41, 96, 99, 103,177–184], tension [178, 185] and gravity loading [46, 186]. The roleof adipose tissue should be considered when in vivo measurementsfor skin are conducted, especially when compression loading isused [12]. However, the contribution of the adipose tissue to themechanical response of the skin might be negligible under nega-tive pressure when the diameter of the aperture, in contact withthe skin, is small (<6 mm) [5, 27]. It has been reported that adi-pose tissue experiences larger strains than the dermis in situationsof negative pressure [27, 63].


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Adipose tissue has been modeled as possessing hyperelastic [10,27, 31, 34, 45, 47, 143, 181, 182, 187, 188], viscoelastic [1, 11, 12, 41, 42,189, 190] and poroelastic [15] material properties. The values givenfor the elastic modulus of subcutaneous adipose tissue are in therange 0.3 - 22.6 kPa [93, 99, 180, 191], that of breast tissue 0.2 - 23.5kPa [20, 33, 128, 189, 192] and that of heel fat pad vary in the range24 - 180 kPa [177, 184].

3.3 MUSCLE

Muscles are subject to large deformations in vivo. The stress statepresent in the muscle is the result of passive and active contribu-tions [21]. Differences in muscles occur between genders [115]. Amuscle exhibits fibre-reinforced, nonlinear, viscoelastic, nearly in-compressible and anisotropic material characteristics [1, 10, 14, 21,53, 79, 119, 123]. During compression, heart muscle might be closeto exhibiting isotropic behaviour [193]. Muscle has been reportedto have linear characteristics only up to 0.3 % rotational strain [194].

The existing experimental data is mostly limited to tensile lon-gitudinal behaviour and force generation during contraction [21,51, 193]. There is little data available on the mechanical charac-teristics of passive skeletal muscle. Passive transverse mechanicalproperties of skeletal muscle of rat have been studied with in vivocompression [195] and in vitro rotation experiments [194]. The me-chanical characteristics of muscle have been evaluated under dif-ferent conditions, i.e., compression [1, 10, 13, 14, 19, 21, 24, 118, 123,181,193,195,196], tension [10,197–199] and shear [194,200,201]. Butalso, magnetic resonance elastography [202], ultrasound [115] andsonoelastography [203] have been applied.

Muscle has been modeled as being hyperelastic [4, 10, 34, 41,49, 181], viscoelastic [1, 13, 19, 51, 119, 204], poroelastic [205], fibre-reinforced [48, 50, 53, 206] and anisotropic [14, 21, 52] material. Theelastic modulus of skeletal muscle at rest is 7 - 790 kPa [10, 24,118,191,196,197,203,207], cardiac muscle around 100 kPa [118] andsmooth muscle 1.4 - 6.8 kPa [118]. Skeletal muscle contraction can


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increase its modulus 8-fold [196, 203] up to 2800 kPa [115]. Due tothe fibrous structure of skeletal muscle, the elastic modulus in par-allel to the fibres can be higher than 14-fold to the correspondingvalue in the perpendicular direction [207].

3.4 CHANGES IN SOFT TISSUE CHARACTERISTICS

The stiffness of soft tissues can vary substantially and influencetheir function [29]. Alterations in tissue material properties are of-ten indications of diseases and pathologies [5, 29]. Therefore, es-timates of soft tissue material properties are important in predict-ing the effect of trauma, evaluating the pathology and the need fortreatment [5, 16, 202].

Soft tissue characteristics and the volume or circumference alterin many neurological or lymphostatic disorders and venous dis-eases [16]. Tumors are often stiffer than the surrounding regionbut their viscosity might not change [103]. The skin is highly vari-able and sensitive to environmental conditions [65]. The biome-chanical characteristics of skin are known to vary significantly withage [3, 134, 161]. Furthermore, the loss in skin firmness becomessignificant after the age of 30 [134, 208]. Since adipose tissue iswell vascularized, it becomes susceptible to ischemia and hypoxia,which influence its mechanical response [96]. Inflammatory pro-cesses associated with breast diseases promote characteristic reac-tions that stiffen tissues: swelling (edema, oedema) and other fi-brotic responses [103]. Soft tissues may swell pathologically dueto impaired lymph flow, a characteristic condition often related topost-surgical or post-traumatic oedema and lymphoedema. In theskin and the adipose tissue, oedema caused by a short term ther-mal injury can increase the water content by 5 and 80 %, respec-tively [209]. Furthermore, lymphoedema can increase the forearmvolume by 44 %, due to excess fluid mostly located in the adiposetissue [210]. In addition, adipose tissue is susceptible to panniculitisand breast tissue to breast cancer, both conditions affect tissue stiff-ness [66, 192, 211]. There are several diseases, which either directly


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(dystrophy) or indirectly (stroke, arthritis and chronic obstructivepulmonary disease), influence the structure of muscle tissue [202].Furthermore, immobilization for five days can induce a 30 % reduc-tion in the muscle mass in the rat limb [212]. Spastic muscle cellshave the tensional elastic modulus that is two times higher thanthat of normal muscle cells [197]. In addition, muscle stiffens signif-icantly (1.6-fold increase in modulus) after long-term compressioninjury [10].

In clinical practice, soft tissues are often evaluated by manualpalpation, which is a subjective method hoping to assess biome-chanical characteristics [16, 23]. However, palpation is impreciseand prone to errors and cannot provide a quantitative assessment.Palpation is mostly used due to the fact that there are no easy-to-use quantitative measurement devices [23, 202]. Several attemptshave been made to design an objective device to analyse soft tis-sues [16, 33, 126] in order to complement medical imaging modali-ties [103, 202]. In vivo measurement of tissue properties could be avaluable addition in the diagnosis of pathology and in monitoringthe development of healing [156, 202]. Many methods in use canreveal particular aspects of disease and injury, but none is capableof providing a complete perspective the state of the tissue [202].


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4 Constitutive modeling ofsoft tissues

A constitutive equation provides a mathematical description of amaterial response to mechanical loading. The aim of constitutivemodeling in biomechanics is to acquire improved knowledge of bi-ological tissues under various loading conditions. Often, constitu-tive models represents an adequate imitation of the behaviour ofreal materials only when specific conditions are met, e.g., specifictemperature, certain strain level, strict deformation rate or the re-quirement of a frictionless system. A simple model might be morerestricted but it is defined with less parameters and therefore itmay be more straightforward to use. On the other hand, a com-plex model might produce the most accurate simulation outcomefor most mechanical problems, but is defined with multiple pa-rameters and therefore achieving a unique solution might be timeconsuming or even impossible to attain. Furthermore, the aimsof the modeling determine the complexity needed for the constitu-tive models. Different clinical applications demand distinct models,e.g., the devices used for clinical application and tissue engineeringdiffer greatly. Therefore, the purpose of the modeling should bewell defined in order to determine and justify the boundaries ofthe model. Classically continuum mechanics has been divided intothree main categories: solid mechanics, fluid mechanics and rheol-ogy. Solid mechanics studies solid objects, fluid mechanics fluidsand rheology evaluates materials with solid and fluid characteris-tics. However, in biomechanics these disciplines are fully integratedas most biological materials are porous and viscoelastic. A short re-view of the selected models used in this thesis is presented in thefollowing section.


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4.1 ELASTICITY

An ideal linear elastic solid follows Hooke’s law: stress is linearlyproportional to strain [213]. The stress response of an elastic mate-rial is time-independent. An isotropic Hookean solid may be illus-trated by a spring and Hooke’s law can be written in the form

F = kHx, (4.1)

where F is force, kH is spring constant and x is deformation. Thiscan also be written as

σ = Eε, (4.2)

where σ is stress, E is elastic modulus and ε is the strain [79, 213].Shear modulus is a function of G = E

2(1+ν), where ν is Poisson’s

ratio. Therefore, for incompressible (Poisson’s ratio ν = 0.5) elasticsolid E = 3G.

Most biological tissues exhibit nonlinear elasticity (hyperelastic-ity) [1]. Therefore, linear elastic model provides often inaccuratepredictions when a material is under a high strain level. The ma-terial might be relatively soft at low strains but it becomes dramat-ically stiffer at high strains, or vice versa. Hence, multiple mod-els have been designed to mimic this nonlinearity. The nonlinearbehaviour of material can be described in several ways, e.g., withgeneral polynomial hyperelastic material model [214], its simplifiedvariants (neo-Hookean [215], Mooney-Rivlin [216] and Yeoh [217]),Ogden [218] and Arruda-Boyce [219] models. The neo-Hookeanhyperelastic material model is the simplest hyperelastic model andcan be used to mimic the nonlinearity of biological soft tissues un-der short-term loading [12, 17, 19, 34, 45, 72]. The strain energy po-tential function of the neo-Hookean model is expressed in the form

U = C10(I1 − 3) +1

D1(Jel − 1)2, (4.3)

where I1 is the first deviatoric strain invariant, Jel the elastic volumeratio and C10 and D1 are parameters defining the shear behaviour

C10 =G0

2=

E4(1 + v)

(4.4)


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Constitutive modeling of soft tissues

and the bulk behavior

D1 =2B0

=6(1− 2v)

E. (4.5)

In equations (4.4) and (4.5), G0 is the initial shear modulus andB0 the initial bulk modulus. Therefore, both C10 and D1 can beexpressed with E and ν. For an incompressible material, the latterpart of equation (4.3) is reduced. As the neo-Hookean materialis defined with only two parameters, optimization of the materialparameters is straightforward.

4.2 VISCOELASTICITY

Viscosity is a measure of a fluid’s resistance to flow and describesthe internal friction of a moving fluid. A fluid with low viscosityflows easily whereas fluid with high viscosity flows slowly. How-ever, in solid mechanics, viscosity describes the time-dependent be-haviour of a solid or porous (solid+fluid) material. An ideal vis-cous material obeys Newton’s law: stress is proportional to rate ofchange of strain with time [213]. Newton’s law of viscosity may bewritten in the form

F = c(dx

dt

), (4.6)

where c is a viscous damping coefficient. Newtonian viscous be-haviour is often illustrated with a viscous dashpot element. Theequation (4.6) may be written as

σ = η(dε

dt

), (4.7)

where σ is tensile stress, η is the Newtonian viscosity coefficientand ε is tensile strain. For an incompressible fluid, η = 3ηG, whereηG is shear viscosity coefficient.

A viscoelastic material has both elastic and viscous character-istics. Hence, the stress response of viscoelastic material is time-dependent. Viscoelastic behaviour of materials is often determined


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Figure 4.1: Stress-strain responses in typical mechanical protocols (stress relaxation, creepand hysteresis) to study viscoelastic materials.

Figure 4.2: Hooke’s spring (E1, E2) and Newton’s dashpot (η) elements used in viscoelasticstandard linear solid (SLS) body.

with stress relaxation, creep and hysteresis tests (Figure 4.1). Vis-coelastic response can be described with spring and dashpot ele-ments arranged in parallel and in series to create a viscoelastic body(Figure 4.2). Viscoelastic models of an infinite number of springsand dashpots can be roughly categorized into two main groups:models with continuous relaxation spectrum and models of fibrousnetwork [220, 221].

Another approach to describe the viscoelastic response of a ma-terial is to use the Prony series of linear equations [222]. The Pronyseries expansion can be used to approximate the shear stress relax-ation response of material and can be written in the form

G(t) = G∞ +n

∑i=1

Gie−t/τi , (4.8)


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where G∞ is long-term shear modulus, Gi shear modulus of the i:thterm and τi relaxation time. It can also be seen that

G(0) = G0 = G∞ +n

∑i=1

Gi, (4.9)

where G0 is the instantaneous shear modulus. Equation (4.9) canbe used to derive another form of equation (4.8)

G(t) = G0 −n

∑i=1

Gi(1− e−t/τi). (4.10)

Furthermore, gi = Gi/G0 is the relative shear relaxation modu-lus [223]. The variables of the Prony series (G∞, Gi, τi) can be de-termined from relaxation and creep data and be implemented intoFE modeling. However, there is no closed form solution for thevariables of the Prony series of the creep response. In addition, inthe FE model, each term of the Prony series increases the numberof global variables and the complexity of the model. Therefore, theuse of a short Prony series is often essential to attain unique resultsfor optimized parameters.

4.3 POROELASTICITY

A poroelastic material has both elastic and porous material charac-teristics. A poroelastic material is usually also viscoelastic, due tothe flow of fluid through the porous medium. Permeability k is ameasure of how well the material allows fluid to flow through it.Permeability is affected by several factors, i.e., the size of the pores,the interconnectibility of the pores, the fluid in the pores (viscos-ity η, density ρ) and the saturation level of the fluid in the pores.Generally, fluid flows easily in a porous material with large andwell-interconnected pores. Permeability can be written in the form

k =νk

gK, (4.11)

where g is gravity, K hydraulic conductivity and νk = η/ρ kine-matic viscosity of the wetting liquid. The fluid flow in saturated


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porous medium can also be described with isotropic permeability

ki =Ks

1 + β√

vw · vwks, (4.12)

where Ks is the dependence of permeability on saturation of thewetting liquid, β is the velocity coefficient, vw is the fluid velocityand ks is the fully saturated permeability. The void ratio e is thefraction of a volume of the pores Vp in the material with respect tothe volume of solids Vs

e =Vp

Vs=

n1− n

, (4.13)

where n ∈ [0, 1] is the material porosity. In the biphasic theory, thetotal stress σt can be expressed with solid σs and fluid σ f stresstensors

σt = σs + σ f = σE − pI, (4.14)

where σE is effective solid stress tensor, p is fluid pressure and I isthe unit tensor.

4.4 FIBRIL-REINFORCEMENT

The fibrous structure can affect the mechanical characteristics ofsoft tissues, especially when they are subjected to large deforma-tions and a high strain rate. Soft tissues are often isotropic or trans-versely isotropic due to fibres [2, 44, 224]. Quantitative structuraldata on the orientation and concentration of collagen fibres can beimportant in describing the behaviour of fibrous soft tissues accu-rately [44]. However, often structural information is not available.The spatial distribution of collagen in skin may be determined withpolarized light microscopy or confocal laser scanning microscopy(CLSM) [225, 226]. Occasionally, the fibres are omitted when theirimplementation has been considered as being unnecessary due totheir small deformations, minor impact on the studied characteris-tics or excessive computational cost [227].


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Figure 4.3: Typical nonlinear tension response of collagen fibres and implementation of anonlinear spring. ε0 toe limit strain.

Multiple models for fibres (collagen, elastin) and fibril-reinforcedsoft tissues (artery, cartilage, muscle, skin, tendon) have been de-veloped [29, 53, 78, 171, 204, 220, 223, 228–235]. The total stress σt offibril-reinforced biphasic material can be written as

σt = σnonfibr + σfibr − pI, (4.15)

where σnonfibr is the stress tensor of the nonfibrillar matrix and σfibr

describes the stress tensor of the fibres. Spring elements can beused to create fibril-reinforced material (Figure 4.3). Often, nonlin-ear springs are set to induce small or zero stress until one obtainsthe toe limit strain ε0 which is then followed by a rapid stiffening.Furthermore, the fibres are often designed to resist only tensionwhich eliminates their role to imitate the skeletal muscle duringcompression [21]. Springs and dashpots in combination can be usedto model viscoelastic fibres [236].


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5 Aims of the thesis

Several pathological conditions induce variations in the mechanicalproperties of soft tissues. Sensitive and objective instrumentation tomeasure tissue stiffness may help to diagnose and monitor the stateof pathologies, e.g. lymphoedema, and aid curative treatments.

The present study aimed to optimize and validate the use ofhand-held devices, based on the application of compression (stiff-ness meter) or external negative pressure (suction), to detect or in-duce changes in soft tissue stiffness.

The specific aims of this thesis were:

1. To investigate the sensitivity of the instrumentation to moni-tor changes in the soft tissue (skin, adipose tissue and muscle)stiffness.

2. To evaluate the effects of soft tissues thicknesses on the sensi-tivity of the stiffness meter.

3. To clarify the effect of physical loading and tissue swelling, asinduced by venous occlusion, on the stiffness of soft tissues.

4. To improve the sensitivity of the stiffness meter by optimizingthe instrument’s geometry.

5. To compare different negative pressure treatment protocolsand to clarify how these protocols affect the interstitial fluidflow in porous soft tissues.


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6 Materials and methods

In four separate studies (study I-IV), the mechanical characteris-tics of human forearm soft tissues (skin, adipose tissue and mus-cle) were studied using experimental and computational methods.Healthy volunteers were recruited for these studies. Two mechani-cal measurement devices, based on compression and negative pres-sure of soft tissue, were used to collect experimental data. Tissuegeometry, implemented in the layered FE models, was determinedutilizing ultrasound (US) and peripheral quantitative computed to-mography (pQCT) imaging. Simulated mechanical behavior of theFE models was matched with the measured responses. The mod-els with optimized material parameters were used to study effectsof the changes in tissue stiffness on the simulated response. Fur-thermore, influences of the changes in the device geometry and thetissue thickness on the simulated response was quantified. Also,two negative pressure treatment methods were also compared andthe impact on intrinsic fluid flow was evaluated.

6.1 SUBJECTS

Nine (seven males, two females) healthy volunteers participated instudy I. Their physical characteristics were as follows (mean ± SD):age = 31 ± 8 years, height = 180 ± 10 cm, weight = 84 ± 10 kg andlength of ulna bone = 29 ± 2 cm. Seven subjects were right-handedand two left-handed. In studies II-IV, eleven (nine males, two fe-males) healthy volunteers were recruited. Their physical character-istics were as follows (mean ± SD): age = 31 ± 7 years, height =179 ± 9 cm, weight = 84 ± 9 kg and length of ulna bone = 29 ±2 cm). Nine subjects were right-handed and two left-handed. Allsubjects gave their informed consent. The studies were approvedby the Kuopio University Hospital Committee on Research Ethics(diary number 7/2010).

The mechanical and imaging measurements were performed on


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Figure 6.1: Illustration of a forearm. The location of the measurements is highlighted(red dot). Humerus (H), ulna (U) and radius (R) bones are simplified and may not beanatomically accurate.

the dominant arm of the subjects. The measurement site was local-ized in the mid ulna (proximal-distal direction), between the ulnaand the radius on the dorsal forearm (Figure 6.1).

6.2 ULTRASOUND MEASUREMENTS

In study I, US (ProSound ALPHA 10, Aloka Co Ltd, Tokyo, Japan)B-mode imaging (5 MHz, a linear array probe (UST-5411)) was con-ducted at the site of mechanical testing, i.e., dorsal forearm (Figure6.2). Three images were taken from each subject. The image pixelsize was 0.051 x 0.051 mm. After calibration, the thickness of dif-ferent tissue layers was analysed from each ultrasound image usingMatlab (versions 7.8.0 - 7.12.0, Mathworks Inc., Natick, MA, USA).Finally, values of tissue thickness of nine subjects were averaged tocreate a single, representative model geometry. Thickness values oftissue layers were (mean ± SD): 2.1 ± 0.4 mm (range 1.7 - 2.9 mm),2.1 ± 2.4 mm (range 0.3 - 8.2 mm) and 10.3 ± 2.0 mm (range 6.4 -13.2 mm) for skin, adipose tissue and muscle, respectively.

6.3 pQCT AND SEGMENTATION

In studies II-IV, forearm was imaged using pQCT at the sites of me-chanical testing (Stratec XCT 2000, Stratec Medizintechnik GmbH,


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Materials and methods

Figure 6.2: Ultrasound image of the forearm of three subjects. Pixel size was 0.051mm x0.051mm.

Pforzheim, Germany). One image slice (thickness = 2.3 mm, pixelsize = 0.2mm x 0.2mm, acceleration voltage 58 kV and current 130µA) was taken for each subject. Different tissue layers were man-ually segmented from the images, and the tissue areas were calcu-lated using Matlab.

For modeling purposes, each tissue cross-section was simpli-fied to consider it as a circular shape (Figure 6.3). In studies II-IV,mean radii (circular approximation) of skin, adipose tissue, muscle,radius (bone) and ulna (bone) were 40.9, 39.3, 36.2, 6.8 and 7.1 mm,respectively. Additionally, in studies II and III, the radii of tissuesfor subject 1 were 38.9, 37.2, 34.7, 7.0 and 7.3 mm and for subject 2they were 40.0, 38.3, 34.8, 7.0 and 7.2 mm, respectively.

6.4 MECHANICAL MEASUREMENTS

In this study, both compression and negative pressure based mea-surement devices were used. The subject’s arm lay on a mechanicalsupport during each test protocol. During the measurements con-ducted at rest (FR: forearm at rest), the subject avoided any mus-cle activity. After a 5 minute break, three isometric (muscle is notallowed to shorten [124]) maximal voluntary contractions (MVC)were performed and the maximum loading of the forearm was de-


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Figure 6.3: (a) pQCT image of the forearm of a subject. Pixel size was 0.2 mm x 0.2 mm.(b) Image after tissue segmentations with arm support removed. (c) Borderlines of tissuesafter segmentation and circular approximation. Difference is highlighted (gray area). (d)Simplified 3D geometry used for FE simulations.

termined using a dynamometer (accuracy ±5 N) (Digitest Force,Digitest PLC, Muurame, Finland). Measurements during isometricflexor loading (IFL, the flexors were contracted) were conducted at50 % load level of the MVC. A short rest was applied before themeasurements were repeated. After another 5 minute break, themeasurements during isometric extensor loading (IEL, the exten-sors were contracted) were performed similarly as the IFL measure-ments. The muscle contractions for IFL and IEL were obtained bypressing the palm downwards or raising it upwards, respectively.After a 5 minute break, venous occlusion (VO), created with a pres-sure cuff at 8 kPa (60 mmHg) for 12 min, induced swelling of theforearm soft tissues. The repeated measurements were conductedat 4, 8 and 12 minutes. During each protocol, 20 repeated mea-


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Table 6.1: Measurement protocols (loading conditions) in each study I-IV. Forearm at rest(FR), isometric flexor loading (IFL), isometric extensor loading (IEL), venous occlusion(VO).

loading Study I Study II Study III Study IV(indentation) (suction) (suction) (indentation)

FR x x x xIFL x - - -IEL x - - -VO x x - -

surements were conducted. The mean response was calculated as areference for model responses. In addition, the mean mechanical re-sponse of the tissues under the VO setup was observed to be nearlyidentical between each time interval (4, 8, and 12 min). Therefore,the VO results were averaged and one response curve was createdfrom the three VO curves. The conducted measurement protocolfor each study I-IV is presented in Table 6.1.

6.4.1 Compression

In studies I and IV, the measurement device was a stiffness (com-pliance) meter (Figure 6.4), a small hand-held indentation devicefor measuring the transient mechanical response of soft tissues. Theforce by which the tissue resists the constant deformation is an indi-cator of the tissue stiffness [16]. It consists of two coaxial load cells:1) 10 kg (Honeywell Sensotec load cell 31/1430-04 Mid, Columbus,Ohio, USA) which measures the total compressing force, and 2)1000 g (Honeywell Sensotec load cell 31/1426-02 Mid, Columbus,Ohio, USA) which measures the indenter force. The sum of thereference plate and indenter forces indicates the total compressingforce. The outer and inner rod diameters were 30.5 mm and 6.6mm, respectively. The indenter diameter and length were 2.0 mmand 1.0 mm, respectively. The instrument was connected to a PCcomputer which was used for the data acquisition using Labview


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Figure 6.4: Schematic presentation of the soft tissue stiffness meter used in studies I andIV. Image modified and reproduced from [16] with kind permission of IPEM by IOPPublishing Ltd ( c©Institute of Physics and Engineering in Medicine, all rights reserved).

software. The sampling frequency for the data acquisition was 1000Hz. The data was recorded through the entire measurement pe-riod. After measurements, the measurement data was analysed us-ing Matlab. During the mechanical compression measurements, thestiffness meter was pressed instantaneously against the skin surfaceuntil a constant pre-defined load (4 N) has been reached.

6.4.2 Negative pressure

The negative pressure (suction) treatment apparatus in use (Lym-phaTouch, HLD Healthy Life Devices, Espoo, Finland) for studiesII and III is a hand-held device for treatment of post-surgical orpost-traumatic oedema and lymphoedema (Figure 6.5). It has aninfrared light sensor which also measures the skin rise (stretch ordeformation) caused by the negative suction pressure. Due to theshape of the forearm, an ellipse shaped suction head was used. Theminor and major axes of the suction head in use are 28.0 and 43.5mm, respectively. During the treatment, the suction head, made of


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Figure 6.5: The negative pressure treatment device used for mechanical measurements andfour replaceable suction heads. The leftmost suction head (attached to the device) was usedduring the measurements conducted in studies II and III.

thermoplastic elastomer SEBS (Thermolast K, Kraiburg TPE GmbH& Co. KG, Waldkraiburg, Germany), makes a tight contact with theskin. The pressure and skin deformation were measured throughthe experiment with accuracies of ±0.01 mmHg and ±10 µm, re-spectively, at a sampling frequency of 40 Hz. The registered datawas stored and analysed offline using Matlab (versions 7.8.0 - 7.12.0,Mathworks Inc., Natick, MA, USA). During the mechanical nega-tive pressure measurements for studies II and III, the treatmentdevice was held on the skin surface until a pre-defined suction pres-sure 13.3 kPa (100 mmHg) was reached.

6.5 FE MODELS FOR SKIN, ADIPOSE TISSUE AND MUSCLE

6.5.1 Model creation and geometry

The FE models were created using Solidworks 2010 and simula-tions were conducted using ABAQUS 6.8 - 6.12 (SIMULIA, Das-sault Systèmes, Providence, RI, USA). In study I, one axisymmetricFE model was created (Figure 6.6). The thickness of each tissue


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Figure 6.6: FE models for studies I (left) and II-IV (right). Axisymmetric model (SA -symmetry axis) was used for stiffness meter simulations and 3D model for simulations of(a) stiffness meter and (b) negative pressure treatment device. Different tissue areas (skin,adipose tissue, muscle) are indicated with different colours.

layer was constant, based on the typical ("mean") thickness valuesof the US measurements. The modeled tissue block had a radius of15 cm. In studies II-IV, one 3D FE model was created (Figure 6.6)based on the mean experimental results, therefore it represented thegeneral properties of subjects. In studies II and III, two additionalmodels were created based on measurement results of two individ-uals with clearly different (”extreme”) tissue responses in the studypopulation. In studies II-IV, the model geometry based on detailedgeometrical pQCT scans of each individual volunteer led initiallyto extremely complex model geometry with very small elementsin some specific regions, and subsequently, high element countand time-consuming computation. Therefore, each modeled tissuecross-section in pQCT was simplified to be of a circular shape. ThepQCT images were used to construct the model geometry by firstsimplifying each tissue cross-section into a circular shape. Then, acircular shaped cross-section area, defined by the circle radius, ofeach tissue in the models was matched with the measured cross-section area. Finally, the 2D geometry was converted to 3D, pro-ducing a cylindrical shape for the model geometry (Figure 6.3). Inthe 3D models, the forearm was 30 cm in length, correspondingwell with the experimentally measured average ulna length. The


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geometries of the simulated devices were based on technical draw-ings and dimensional measurements taken with a digital caliper(accuracy ±0.01 mm).

6.5.2 Elements and boundary conditions

In all studies, optimal density for the FE mesh was found afterrunning a convergence study [227]. The mesh at the areas of in-terest (usually the skin and in close proximity of the indentationand pressurization area) was made denser in the analyses to ob-tain smoother distributions of the model output. In all models, thecontact between the meter and the skin layer was assumed to befrictionless. In study I, the stiffness meter consisted of 3990 (CAX3and CAX4R) and the soft tissues of 11000 axisymmetric elements(CAX4R). The vertical movement of the bottom nodes of the mus-cle layer was restricted. In study II, the suction head consisted of3100 - 6800 rigid 4-node surface elements (R3D4), the tissues of24900 - 33000 hexahedral 8-node elements with reduced integrationpoints (C3D8R) and the skin fibres of 18600 - 22900 spring elements(SPRINGA). The ulna and radius bones in the model were consid-ered to be rigid and static by restricting the movement of the mus-cle nodes in contact with the bones. Both ends of the muscle tissuewere fixed. As the perimeter of the bones was fixed to be station-ary, there was no need to use elements inside the bones. In studyIII, on average, the suction head consisted of 700 rigid 4-node sur-face elements (R3D4), the adipose tissue of 11800 hexahedral 8-nodeelements with reduced integration points (C3D8R) and the skin fi-bres of 47600 spring elements (SPRINGA). The skin and the muscleconsisted of 17700 and 7900 hexahedral 8-node pore pressure ele-ments with reduced integration (C3D8RP), respectively. Fluid flowthrough the skin was prohibited but it was allowed through theend regions of the skin and the muscle. In study IV, on average,the skin fibres consisted of 41400 spring elements (SPRINGA) andthe device and the tissues consisted of 1900 and 31600 8-node 3Delements with reduced integration (C3D8R), respectively.


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Figure 6.7: Schematic presentation of simulated cyclic and continuous negative pressuretreatment methods. The location of the suction head along z-axis (top) and the inducednegative pressure (bottom) are presented. (a) Pressurization phase, (b) movement phase,(c) end of last pressurization phase and (d) beginning of relaxation phase (at the sametime the suction head is lifted from the skin) for the cyclic negative pressure method. (e)Pressurization phase, (f) movement (and pressurization) phase, (g) end of movement phaseand (h) beginning of relaxation phase for the continuous negative pressure method.

6.5.3 Loading protocol

In studies I and IV, the simulated compliance meter was pressedagainst the forearm with the same force as used in the experi-ments. In studies II-III, the negative pressure field was modeledto be shaped as a portion of a sphere [26], exposed on the skin sur-face within the pressurization area. Additionally, in study III, porepressure and fluid velocity under two different negative pressuretreatment protocols, postulated to have clinical utility, were simu-lated. During both methods, soft tissues were loaded under nega-tive pressure for 5 s. The first method consisted of repeated cyclicnegative pressures with a change in location between each pressur-ization, and the second of continuous pressurization while movingthe suction head with a constant speed (Figure 6.7). The simulated


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Table 6.2: Material models for each soft tissue in studies I-IV. Hyperelastic (HE), fibril-reinforced hyperelastic (FRHE), fibril-reinforced porohyperelastic (FRPHE), viscohypere-lastic (VHE), porohyperelastic (PHE).

tissue Study I Study II Study III Study IVskin HE FRHE FRPHE FRHE

adipose tissue HE HE VHE HEmuscle HE HE PHE HE

cyclic procedure consisted of five 13.3 kPa pressurizations and 1cm lengthwise movements (towards treatment direction) betweeneach pressurization (4 cm total movement). The maximum nega-tive pressure was applied within 1 s, and subsequently removedwithin 0.01 s. There was a 1 s zero pressure period, consisting oflift, movement to a new location and a new skin contact of the suc-tion head, between each pressurization. The continuous treatmentprotocol consisted of one 13.3 kPa pressurization (within 1 s) withsimultaneous 4 cm (in 4 s) lengthwise movement (towards treat-ment direction) at a constant speed.

6.5.4 Material model description

In all studies, neo-Hookean hyperelastic (HE) material model wasused for all soft tissues to imitate the characteristic nonlinearity ofbiological soft tissues (skin, adipose tissue, muscle) under short-term loading. Additionally, in studies II and IV, the model of theskin was expanded to be fibril-reinforced hyperelastic (FRHE). Fur-ther, in study III, the skin was modeled to be fibril-reinforced poro-hyperelastic (FRPHE) [32, 41, 129], the adipose tissue viscohypere-lastic (VHE) [1,12] and the muscle as being porohyperelastic (PHE)[181,205] (table 6.2). The Poisson’s ratio for each tissue was fixed to0.4, similarly to previous studies [10,21,26,30,128,146,170,191]. Theelastic modulus of the muscle was fixed to 100 kPa [10,24,203,237],except in study I for the IEL, model the elastic modulus of the con-tracted muscle was fixed to 400 kPa [24, 200, 203]. In studies II-IV,


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fibril-reinforcement was added to the skin to mimic the mechanicalbehavior of the natural skin more accurately. This was realized byusing nonlinear springs (SPRINGA element type in Abaqus), whichwere aligned lengthwise along the tension axes. The properties ofthe collagen fibres could then be described by the fibril networkmodulus E f = ∆σ/∆ε (describing the rigidity of straightened fib-rils). The fibril network modulus of the skin was fixed to a value of4.4 GPa [64, 175]. Toe limit strain ε0 was also added into the modelto describe the possible effect of fibril straightening [5, 238, 239];larger ε0 indicates longer tissue elongation before collagen fibreshave been straightened. Study II revealed that the FRHE materialwas able to capture the mechanical tensional response of the skinpresented experimentally in [154]. The viscoelastic component ofthe adipose tissue (in study III) was described by using the Pronyseries expansion for shear relaxation function. Viscoelastic param-eters for adipose tissue were n = 1, g1 = 0.39 and τ1 = 700 s [12],where n is the number of terms in use, g1 the relative shear relax-ation modulus and τ1 relaxation time. The fluid flow in saturatedporous medium (skin, muscle) in study III was described by usingisotropic permeability. For simulation in study III, three permeabil-ity related parameters were defined as follows: γw is specific weightof the wetting liquid (assumed to be that of water, 9820 N/m3), Khydraulic conductivity and e void ratio. The hydraulic conductivityvalues for skin and muscle, 10−6 and 10−7 m/s, respectively, wereused [240, 241]. Furthermore, the void ratios for skin and musclewere 0.86 and 0.13, respectively [242].

6.5.5 Sensitivity analysis and optimization

First, the importance of the tissue parameters for the simulated re-sponses was studied by conducting preliminary sensitivity analysis(explained in more detail below). Between two to three of the mostinfluential tissue parameters were optimized, and the rest werefixed as constant. In study I, optimization of the material propertieswas conducted for the elastic modulus of the skin and the adiposetissue, and in studies II-IV, also for the toe limit strain of the skin


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fibrils. The material parameters were optimized by minimizing themean squared error (MSE) between the simulated and experimentalresponses using a multidimensional unconstrained nonlinear min-imization routine (fminsearch) available in Matlab. MSE is definedas follows:

MSE =1n

n

∑i=1

(Rmodel,i − Rexp,i)2, (6.1)

where Rmodel,i and Rexp,i are the response values obtained from themodel and experimental data (in studies I and IV indenter forceand in studies II and III tissue deformation). After parameter op-timization, the sensitivity of the simulated responses to changes intissue parameters and model geometries was analyzed (table 6.3).The elastic modulus values of the soft tissues (skin, adipose tissueand muscle) and the toe limit strain of the skin fibrils were changedindependently, usually from 50 % to 200 % of the optimized, ref-erence values, and their effects on the mechanical response of theforearm were analyzed. In study IV, an additional sensitivity anal-ysis was done for twelve different indenter geometries and the skinand adipose tissue volumes (∼thickness) were changed from 50 to200 % of the measured. The studied indenter geometries were asfollows: indenter width 2 and 2.5 mm; length 0.5, 1.0 and 2.0 mm;radius of the rounding (head shape) <1, 1 and >1 mm. The in-denter head in the simulations was constructed by using differentrounding geometries: the first geometry consisted of filleted edges

Table 6.3: The sensitivity of the simulated responses to changes in tissue parameters andmodel geometries analyzed in studies I-IV.

Change in model Study I Study II Study III Study IVelastic modulus x x x x

fibres - x x xviscosity - - x -porosity - - x -

tissue thickness - - - xdevice geometry - - - x


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with radii of 0.1, 0.2 and 0.4 mm (termed “corner”); the secondgeometry consisted of a sphere using a radius of 1 mm (termed“sphere”); the third geometry consisted of a rounded indenter tipthat matched the curvature of a sphere with a radius of 1.625 and2.5 mm (termed “oval”).


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7 Results

7.1 INDENTATION

7.1.1 Experimental measurements

In study I, the highest indenter forces in the forearm were measuredduring IEL (N = 9). VO increased slightly the reaction forces, ascompared to those in FR. During IFL, the indenter forces were be-tween those during IEL and VO. During FR, VO, IFL and IEL, theaverage experimental indenter forces were 0.33 ± 0.07 N, 0.39 ±0.08 N, 0.54 ± 0.16 N, 0.66 ± 0.20 N (mean ± SD) under a totalforce of 3.93 N (Figure 7.1), respectively. In study IV, during FR,the average experimental indenter force was 0.32 ± 0.07 N (N = 11)under a total force of 3.93 N.

7.1.2 Simulation

In study I, values of the elastic modulus of skin and adipose tissues,as obtained by fitting the model to the mean experimental mechan-ical responses of the forearm at rest, were 210 kPa and 1.92 kPa,respectively. During IFL, IEL and VO, the optimized value of theskin elastic modulus was 112, 210 and 21 % higher than that at rest,respectively. The elastic modulus of the adipose tissue was 47, 32and 47 % smaller after application of IFL, IEL and VO, respectively,than the values at rest. In study IV, values of the elastic modulus ofskin and adipose tissue, as obtained by fitting the model to the ex-perimental data, were 130 and 2.52 kPa, respectively. The optimizedfibril toe limit strain was 7.19 %.

While the simulated total compressing force for the stiffness me-ter was 2 N in study I, the highest von Mises stress values werepresent in the skin layer, beneath the indenter. The highest strainvalues were concentrated at the interface between the skin andadipose tissue. The maximum von Mises stress value under theFR/VO/IFL/IEL condition was 85.8/95.1/125.6/158.5 kPa for the


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Figure 7.1: The mean experimental indentation results recorded (N = 9) at rest (FR), un-der venous occlusion (VO), isometric flexor loading (IFL) and isometric extensor loading(IEL). Standard deviation is shown shadowed.

skin, 4.6/8.3/8.0/7.3 kPa for the adipose tissue and 3.0/3.9/3.8/3.8kPa for the muscle. Similarly, the maximum logarithmic strainvalue under the FR/VO/IFL/IEL condition was 0.180/0.171/0.139/0.124 for the skin, 0.169/0.166/0.140/0.125 for the adipose tissueand 0.011/0.014/0.013/0.004 for the muscle.

The simulated indenter force was most sensitive to changes inthe elastic modulus of the skin layer, while adipose tissue, muscleand toe limit strain of the skin fibrils (in study IV) had less effectson the indentation response of the forearm. In study I, the variationin the elastic modulus of the skin layer (-50 – 200 % of the referencevalue) changed the indenter force by -33 – 45 %. The correspondingvalues for the effect of the adipose tissue and muscle layer on theindenter forces were -6 – 11 % and -5 – 4 %, respectively. Changesin the modulus of the adipose tissue controlled the slope of the in-denter force versus total force curve. In study IV, the simulatedresponse was 3.9 ± 1.2 (mean ± SD, range 2.4-6.3), 4.9 ± 2.0 (range1.7-8.7) and 2.2 ± 0.7 (range 1.4-3.8) times more sensitive to changesin the elastic modulus of the skin than to those of the adipose tis-sue, muscle and toe limit strain of the skin fibrils, respectively. Theindenters of 0.5 mm in length most sensitively detected the stiffnesschanges in the skin, whereas indenters with lengths of 1 and 2 mm


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Results

Figure 7.2: Sensitivity of the simulated indenter force to the changes in the elastic modulusof skin, adipose and muscle tissues, and the fibril toe limit strain of the skin fibrils. Re-sponses using 50 % or 200 % of the optimized material values were examined to determinesensitivity of the measured indenter force to changes in the values of material parameters.The same sensitivity analyses were conducted for indenters with lengths of 0.5, 1, and 2mm.

were 6.5 - 13.6 and 28.5 - 32.4 % less sensitive, respectively (Figure7.2). Similarly, the indenters of 2 mm in diameter revealed the stiff-ness changes in the skin most sensitively, whereas the indenters of2.5 mm in diameter were 10.9 - 15.9 % less sensitive. Moreover, thespherical-ended indenters detected the stiffness changes in the skinmost sensitively, whereas indenters with oval and corner roundedends were 0.8 - 8.3 and 4.4 - 12.5 % less sensitive, respectively.

In the simulations (study IV), changes (50 and 200 %) in theskin volume had a relatively higher impact on the sensitivity of thestiffness meter to measure the elastic modulus of the adipose tissueand the muscle (±53.1 - 57.4 %), as compared to the modulus ofthe skin and the fibril toe limit strain (±13.5 - 25.6 %). A similarchange in the volume of the adipose tissue had also a relativelymajor effect on the sensitivity of the stiffness meter in measuringthe elastic modulus of the muscle (±69.7 %), as compared to themodulus of the adipose tissue and fibril toe limit strain (±9.1 - 9.2%) and the modulus of the skin (±5.3 %) (Figure 7.3).


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Figure 7.3: Sensitivity of the simulated indenter force to the changes in the elastic modulusof skin, adipose and muscle tissues, and the fibril toe limit strain of the skin fibrils. Re-sponses using 50 % or 200 % of the optimized material values were examined to determinesensitivity of the measured indenter force to changes in the values of material parameters.The same sensitivity analyses were made with the volumes of 50 %, 100 %, and 200 % ofthe measured volume of the skin and adipose tissue.

7.2 NEGATIVE PRESSURE

7.2.1 Experimental measurements

In study II, after the VO, the skin deformation was reduced by 7.9± 9.5 % (N = 11) (Figure 7.4). The measured differences in re-sponses during FR and VO were statistically significant (p < 0.05).In study III, during FR, the average experimental tissue deforma-tion was 4.55 ± 0.71 mm (mean ± SD, range 3.86-5.99 mm, N = 11)under maximal negative suction pressure. Similarly, the deforma-tion under maximal suction pressure was 3.90 ± 0.34 and 5.99 ±0.33 mm for subjects 1 and 2, respectively.

7.2.2 Simulation

In study II, values of the elastic modulus of skin and adipose tissue,as obtained by fitting the model to the experimental data at rest,were 1130 ± 210 kPa and 0.80 ± 0.36 kPa (mean ± SD), respectively.The fibril toe limit strain at rest was 2.78 ± 1.41 %. As a result of theVO, the optimized modulus values of the skin and adipose tissueincreased by 27 ± 21 % and 35 ± 26 %, respectively. In study III, the


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Results

Figure 7.4: Experimental negative pressure results (Mean ± SD) at rest (FR) and undervenous occlusion (VO) for whole subject group (N = 11) and subjects 1 (s1) and 2 (s2).Standard deviation is shown shadowed.

values of the elastic modulus of skin and adipose tissue were 930± 530 kPa (range 440 - 1480 kPa) and 0.84 ± 0.19 kPa (range 0.62 -0.97 kPa), respectively. The optimized value for the fibril toe-limitstrain was 2.22 ± 0.32 % (range 1.96 - 2.58 %) (table 7.1).

In study II, the highest von Mises stress values were locatedin the skin layer, at the longitudinal edges of the instrument’s suc-tion head. The highest strains occurred in the adipose tissue layer,and were located under the center of the suction head. The maxi-mum von Mises stress value under the FR/VO condition was 61.2± 8.9/64.1 ± 13.7 kPa for the skin, 1.0 ± 0.2/0.9 ± 0.2 kPa for

Table 7.1: Simulated optimized tissue parameters in studies I-IV under FR (forearm atrest). Elastic modulus E of the skin and the adipose tissue (ad. tis.), the toe limit strain ofskin fibrils ε0, indentation (ind.).

parameter Study I Study II Study III Study IV(ind.) (suction) (suction) (ind.)

Eskin (kPa) 210 1130 ± 210 930 ± 530 130Ead. tis. (kPa) 1.92 0.80 ± 0.36 0.84 ± 0.19 2.52ε0 (%) - 2.78 ± 1.41 2.22 ± 0.32 7.19


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the adipose tissue and 1.6 ± 0.6/2.0 ± 0.6 kPa for the muscle. Themaximum logarithmic strain value under the FR/VO condition was0.04 ± 0.01/0.04 ± 0.00 for the skin, 0.94 ± 0.07/0.87 ± 0.05 for theadipose tissue and 0.02 ± 0.01/0.02 ± 0.01 for the muscle.

In study II, the sensitivity of the deformation registered by theinstrument to different tissue properties varied along the level oftissue deformation. Under small deformation (<2 mm), the modelresponse was up to 13 % more sensitive to changes in the elasticmodulus of the adipose tissue than to those of skin. When therewas only a large deformation (>3 mm), the model response wassubstantially more (up to 86 %) sensitive to changes in the elasticmodulus of the skin than to those of adipose tissue. Further, withlarge deformations, the modeled response was up to 54 % moresensitive towards the changes in the fibril toe limit strain than inthe elastic modulus of the adipose tissue. The modulus of muscleand the fibril network modulus of skin (when >50 MPa) had onlyminor effects on the model response. In study III, modeled tissuedeformation under negative pressure was most sensitive to changesin the elastic modulus of skin and adipose tissue as well as the fibriltoe limit strain, causing 12.9 - 21.3 % difference in the simulatedresponse after changes in the parameter value. The Poisson’s ratioand the relative shear relaxation modulus of the adipose tissue hadlower (7.6 - 8.4 %) impacts on the simulated response, while therest of the parameters had minor (<1.9 %) effects on the simulatedresponse (table 7.1).

In study III, the simulated pore pressures within the skin tissuelayer during the cyclic treatment were -1.0 – 6.2 and -15.3 – 13.3kPa at the end of pressurization and at the beginning of relaxation,respectively. Similarly, the pore pressures during the continuoustreatment were -2.7 – 6.2 and -23.6 – 15.4 kPa at the end of suctionand at the beginning of relaxation, respectively (Figure 7.5). Af-ter (the last) pressurization and at the beginning of the relaxationphase, global maximum and minimum values of the pore pressurein the skin were reduced to 10 % from the peak value within 0.6 -0.7 and 0.3 - 1.0 s during cyclic and continuous negative pressure


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Table 7.2: Sensitivity of the models to changes in the material parameter values, as indi-cated by the mean difference between the simulated response of the optimized and analysedFE model in studies I-IV. Elastic modulus (E), Poisson’s ratio (v), fibril network modulus(Ef), fibril toe limit strain (ε0), hydraulic conductivity (K), void ratio (e), relative shearrelaxation modulus (g1), relaxation time (τ1), parameter (par.), adipose tissue (ad. tis.),indentation (ind.), negative pressure (n.p.).

par. tissue change in the difference, studyvalue (%) ind./n.p. (%)

E skin 50-200 31.6-37.9/20.9-21.3 I-IVad. tis. 50-200 5.6-9.7/14.1-16.5 I-IVmuscle 50-200 3.2-7.5/1.6-1.9 I-IV

v skin 87.5-112.5 0.5/0.5 I, IIIad. tis. 87.5-112.5 2.6/7.6 I, IIImuscle 87.5-112.5 0.2/0.2 I, III

Ef skin 50-200 -/0.0-0.1 II-IIIε0 skin 50-200 18.4/12.9-14.8 II-IVK skin 10-1000 -/0.3 III

muscle 10-1000 -/0.0 IIIe skin 50-200 -/0.3 III

muscle 50-200 -/0.0 IIIg1 ad. tis. 50-200 -/8.4 IIIτ1 ad. tis. 50-200 -/0.0 III

methods, respectively.

In study III, the maximum simulated fluid velocity in the skinwas 0.44 (cyclic) and 0.46 mm/s (continuous), comprised mostly ofthe y- and z-axis component, with the two different negative pres-sure methods. The maximum simulated fluid velocity in the skinalong the treatment direction of the forearm (along z-axis) was upto 0.25 and 0.37 mm/s during cyclic and continuous pressuriza-tions, respectively. During cyclic pressurization, the mean velocityof the pore fluid in the simulated skin along the treatment directionwas less than 4 % of that found under continuous pressurization(Figure 7.6). Furthermore, the mean fluid velocity during the firstcyclic pressurization was less than 30 % of that found during the fol-


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Figure 7.5: Simulated pore pressure (POR) at the end of suction (a,c) and at the begin-ning of relaxation (b,d) for cyclic and continuous suction methods. Suction proceduresprogressed towards positive z-axis.

lowing pressurizations. The average fluid displacement in the skinalong the treatment direction was 3.4 · 10−6 and 89.9 · 10−6 mm dur-ing simulated cyclic and continuous pressurizations, respectively.


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Figure 7.6: Simulated z-axis component of the mean fluid velocity during cyclic and con-tinuous pressurization methods. (a-e) Pressurizations 1-5 for the cyclic negative pressuremethod. (f) Pressurization phase and (g) movement (and pressurization) phase for thecontinuous negative pressure method.


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8 Discussion

The conventional method, palpation is a subjective technique usedto assess the condition of the tissues but it is both imprecise andprone to errors. It is still in use mostly because no sensitive andsimple experimental instrumentation is available. A sophisticatedin vivo measurement device for soft tissue properties would be valu-able since it could sensitively diagnose pathologies and monitorthe development of healing. Verification and optimization of theexperimental instrumentation and the treatment efficiency of nega-tive pressure therapy necessitates a clearer understanding of the tis-sue straining, fluid pressurization and flow characteristics in fibril-reinforced porous soft tissues.

In this study, experimental measurements of the forearm ofhealthy volunteers were conducted at rest and under loading con-ditions. Geometry of the forearm tissues was measured using USand pQCT imaging. Novel FE models of the forearm were createdand the model responses were matched with those of the exper-imental measurements. Parametric sensitivity analyses were con-ducted by evaluating the impact of the changes on the stiffness andthickness of the tissues and in the device geometry on the modelresponses. Finally, two experimental negative pressure treatmentprotocols (continuous and cyclic) were simulated to evaluate theirimpact on interstitial fluid flow in soft tissues.

8.1 EXPERIMENTAL MEASUREMENTS AND SOFT TISSUEDIAGNOSTICS

During VO, the mean indenter force increased by 15 % and theskin deformation under negative pressure was reduced by 8 % as afunction of time (0 - 12 min). These results are in agreement withan earlier indentation study that reported ∼13 % tissue stiffeningduring 12 min VO [16]. The highest indenter forces in the forearmwere measured during IEL. During IFL, the indenter forces were


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between those during IEL and VO. The measured mean indenterforce was 53 and 91 % higher during IFL and IEL than the valueat rest, respectively. These results suggest that the studied instru-mentation may be used to objectively monitor changes in soft tissuestiffness.

The 8 kPa (60 mmHg) cuff pressure was considered to providesufficient external pressure to ensure that the superficial veins re-mained occluded [243]. The short-term VO prevents the veins fromtransporting the blood towards heart and this causes the veins toswell, observed experimentally in this thesis mostly by local stiffen-ing of tissues. As little as 4 min of VO produced swelling equilib-rium in the soft tissues.

Non-invasive in vivo tests on human subjects has limitations,e.g., due to ethical and time issues. Traditional mechanical in vitrotests would be valuable in helping to understand soft tissues inmore detail than is possible with traditional in vivo tests. With invitro tests, the mathematical equations that describe the materialcould be defined. Then, the in vivo tests conducted in this studycould be used to fine-tune the parameters of the tissues. It is alsodifficult to study extensively the complex mechanical characteristicsof soft tissues in vivo. Therefore, in future studies, a combination ofin vitro tests with in vivo fine-tuning might be worthwhile.

8.2 MODEL OPTIMIZATION

In all studies, values of the the elastic modulus of the skin and theadipose tissue were optimized since they modified sensitively thesimulated responses. Additionally, in studies II-IV, the toe limitstrain of the skin fibrils was optimized. With these parameters, themodels mimicked the experimental responses successfully. Inclu-sion of a higher number of optimized parameters or more complexmaterial models would have compromized the uniqueness of theoptimization solution. Furthermore, the applied material modeland its parameters are appropriate and sufficient for potential clin-ical applications.


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The present values of the elastic modulus for the tissues are inagreement with those reported earlier. Although, the elastic mod-ulus of the skin in simulated conditions of negative pressure areslightly higher (930 - 1130 kPa) than those reported earlier (0.3 -850 kPa). The optimized values of the elastic modulus of the skinare lower for indentation (130 - 210 kPa) than for negative pressureloading (Table 7.1), which has also been noted in earlier reports.Presumably, this difference is caused by the natural pre-tensionqualities of the skin in vivo, which could not be measured in thepresent study. Natural pre-tension of the skin can be considered asone reason that the in vivo response is shifted in comparison withthe ex vivo response. Indeed, the observed in vivo response resultsfrom a coupled effect of the skin under pre-tension and of the ad-ditional loading [244, 245]. Hence, under negative pressure, skinis further stretched beyond the pre-tension value and is perceivedas being stiffer than under indentation. The optimized values forthe elastic modulus of the adipose tissue vary slightly: 1.92 - 2.52kPa under indentation and 0.80 - 0.84 kPa under negative pressure.In addition, the optimized values for the toe limit strain of skinfibrils differ between indentation (2.2 - 2.8 %) and negative pres-sure (7.2 %). Presumably, these differences are also caused by thenatural loading of tissues and the difference in the experimentalloading. Due to the presence of the skin, adipose tissue in vivo isunder a natural pre-compression and is further compressed duringindentation measurements and is, therefore, perceived stiffer thanduring negative pressure loading. In the future, it may be necessaryto measure pre-tension of the skin to obtain uniform results fromindentation and negative pressure loading.

8.3 SENSITIVITY ANALYSIS

8.3.1 Tissue stiffness

In general, the simulated responses were most sensitive to changesin the elastic modulus of the skin and the adipose tissue and thetoe limit strain of the skin fibrils (Table 7.2). With respect to the


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optimized parameters, the elastic modulus of the skin was pre-dominantly dominating the simulated response in all geometriesand loading conditions. However, when there is a small deforma-tion, the model response may also be sensitive to changes in theelastic modulus of the adipose tissue. The Poisson’s ratio, the rela-tive shear relaxation modulus of the adipose tissue and occasionallyalso the elastic modulus of the muscle had a moderate impact onthe simulated response, but the remainder of the parameters ex-erted only a very minor effect on the simulated response. It shouldalso be noted that even though the constant (fixed) parametershad only a minor effect on the simulated suction response (tissuestretch) in study III, the influence on interstitial fluid flow might beof greater importance. Based on these results, both of the studiedhand-held instrumentation could be used to measure the mechan-ical characteristics of human soft tissues, sensitively altered, e.g.,by pathological tissue swelling. Furthermore, these results clearlyhighlight the importance how skin properties can influence the re-sults obtained with the experimental measurement device.

The indentation simulation indicates stiffening of skin and soft-ening of adipose tissue after application of IFL, IEL and VO. Musclecontraction in IFL and IEL changes muscle geometry, and subse-quently skin strain and stiffness may increase locally. A contractedmuscle becomes shorter, thicker and stiffer, but its volume remainsrelatively constant [113, 121, 122]. However, tissue geometry duringa muscle contraction and under VO was not known. Therefore, thechange in the tissue volume during these loading conditions wasnot considered. With short-term VO, the blood vessels of the fore-arm swell, this being observed in the simulations mostly as stiffen-ing of the skin. In addition, under VO, excess fluid may seep intoadipose tissue, leading to tensile straining of the skin layer and asubsequent stiffening of the skin. The potential physiological ex-planations for these changes in the adipose tissue remain open, al-though model simplifications might have contributed to this ratherunforeseen result. There were very little absolute variations in theproperties of adipose tissue between different loading conditions.


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Discussion

As the adipose tissue modifies the mechanical responses less thanthe skin, softening of the adipose tissue seems to be less meaningfulthan stiffening of the skin.

The elastic modulus of adipose tissue significantly affected theslope of the total force vs. indenter force curve. In compressiveindentation testing, the adipose tissue layer is further compressed(past natural pre-compression) between the skin and muscle, andthis probably exerts a greater effect on the mechanical response.The impact of the adipose tissue on the slope of the simulatedresponse was low during negative pressure loading. Apparently,negative pressure loading creates tension in the pre-tensioned skinand pre-compressioned adipose tissue and, therefore, will furtherstiffen the skin and minimize the effect of the adipose tissue.

Variations (50 - 200 % of the reference value) in the values ofthe fibril network modulus had negligible effects on the model re-sponse. Only values lower than 50 MPa increased the error be-tween the model and the experiment. This finding was related tothe rapid stiffening of the collagen after the fibril toe limit strain hasbeen exceeded. In fact, collagen has a tensile strength comparableto steel [246] and the fibrils might swiftly turn into being virtuallyunstretchable (atleast for the forces described in the model).

8.3.2 Tissue thickness

In study IV simulations, the thickness of skin layer strongly af-fected the ability of the indentation device to detect changes in themechanical characteristics of the underlying adipose and muscletissue. Furthermore, thickness of skin layer had a moderate im-pact on the sensitivity of the stiffness meter to measure the skinand the fibril toe limit strain. Similarly, changes in the amount ofadipose tissue affected the device’s sensitivity to detect the proper-ties of muscle tissue. However, the thickness of the adipose tissuehad only a minor impact on the sensitivity to record the stiffness ofoverlying skin. On the other hand, rapid changes in the fat layerthickness (e.g., after losing or gaining weight) may influence theinitial (“relaxed”) pre-tension level of the skin, and thereby modify


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the skin stiffness. Similar analysis for the effect of changes in thethickness of the tissue layers to the negative pressure loading willneed to be conducted in the future.

8.3.3 Stiffness meter geometry

In study IV, the simulation indicated that the stiffness meter wasmore sensitive to record changes in the skin properties when ashorter and spherical-ended indenter was used. However, from apractical point of view, a very short indenter might be prone tomeasurement uncertainties caused by a non-flat skin surface. Inaddition, a short indenter can complicate the measurements con-ducted at a right angle (perpendicular to the skin). In the experi-ments, a 1 mm long indenter was found to be adequately sensitiveand also user-friendly. Longer indenters could detect more sensi-tively the stiffness changes occurring in the adipose and muscle tis-sues. However, the impact of the skin on the response was alwaysdominating. Therefore, it might prove difficult to achieve highlyreliable measurements of the stiffness changes in the adipose orthe muscle tissue. Based on simulations, an indenter with lowerdiameter was also more sensitive at detecting changes in the stiff-ness of skin. A more spherical-ended indenter seemed to improvesensitivity to measure stiffness changes in the skin. At a constantforce, a spherical-ended indenter, as compared with a plane-endedindenter, produced larger strains in the skin layer and this mayhelp to achieve higher sensitivity. The importance of the size of thereference plate on the simulated response was not systematicallyanalyzed and should be examined in the future.

8.4 FLUID FLOW

In study III, there were substantial differences in the pore pres-sure distribution in skin between the cyclic and continuous nega-tive pressure protocol models. During the cyclic pressurization, thepore pressure distribution was almost symmetric during all pres-surization and relaxation phases. On the contrary, the pore pres-


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Discussion

sure distribution was more complex during the continuous nega-tive pressure. During pressurization, the cyclic method provokedlocal overpressure (>0 Pa) around the pressurization area, whilethe continuous method induced both local overpressure and un-derpressure (<0 Pa) in regions both inside and outside of the pres-surization area. The pore pressure distributions during relaxationphases were quite symmetric and similar for both cyclic and con-tinuous suctions, in spite of the slightly larger area experiencingoverpressure and the smaller area exposed to underpressure undercyclic loading.

The local transient fluid velocities were only slightly (3 %) lowerduring the cyclic negative pressure protocol, but the mean fluid ve-locity was significantly (96 %) lower than that obtained with thecontinuous negative pressure protocol. As the locations of cyclicpressurizations were fixed, only a minor total fluid movement alongthe treatment direction took place. During the continuous proto-col, the suction head moves smoothly over the skin surface and thenegative pressure is simultaneously applied. Then, the fluid moveseffectively along the motion path, i.e., treatment direction.

Overall, the continuous pressurization method transports thefree fluid more efficiently. Considering the poroelastic nature of thetissues, the continuous suction might be the preferred technique foreffective treatment. Due to the lack of experimental data on perme-ability, adipose tissue could not be set to be porous. In the future,the poroelastic properties of adipose tissue should be experimen-tally characterized and implemented into the FE model.

The impact of the simulated negative pressure therapies on theflow of interstitial fluid in interstium was studied. However, vas-cular and lymphatic systems were not implemented in the model.Therefore, the impact of the negative pressure therapies on the flowof fluids (blood, interstitial fluid and lymph) in vessels and intersti-tium was not analyzed. In the future, the contributions of vascularand lymphatic systems should be implemented into the FE model.


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8.5 VALIDITY AND LIMITATIONS OF MODELS

Generally, a major aim in simulation is to mimic reality and to studythe characteristics of materials that cannot be measured experimen-tally. This may often lead to overly complex model descriptions andexcessive computation requirement. Therefore, additional aims insimulation are to minimize computation time, complexity of themodel and the number of unknown material parameters. Simplifi-cations often decrease computation time and the complexity of themodel but also introduce some limitations. Thus, models are usu-ally compromises between accuracy and computation time. Pre-sumably, due to these compromises and the complexity of skin,no mechanical model of the skin has been fully accepted by thisfar [158]. Clearly there still are some simplifications and limitationsin the FE models used in the present studies.

It is challenging to achieve a realistic soft tissue geometry in theforearm, partly because the amount and shape of each tissue altersfrom wrist to elbow. In study I, the geometry of the model wassimplified to be axisymmetric. In studies II-IV, the 3D geometry ofeach tissue was simplified and modeled to be of cylindrical shapebecause fully realistic tissue geometry, reproduced from pQCT im-ages, created many areas with distorted elements in the model.These simplifications may slightly change the absolute values ofthe material parameters obtained from the optimization. However,those should not influence the main conclusions of the thesis as thedifference between the measured and the simplified geometry wasusually lowest in the close proximity to the indentation and nega-tive pressure measurements (Figure 6.3c). In addition, subject spe-cific models could slightly modify the parameter values obtainedin these studies. Furthermore, the studies consisted of individualswith different body types, ages and genders, leading to variationsin thickness of tissue layers among subjects.

In study I, skin anisotropy was not implemented into the FEmodel, due to simplicity required by the optimization process toderive a unique set of parameter values. In studies II-IV, skin fibres


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were aligned in parallel to the skin surface, making skin anisotropic.Due to the simplicity required by the optimization process, longitu-dinal and transversal fibres could not be differentiated. In addition,due to the potential clinical application of the model, there was noneed to differentiate longitudinal and transversal fibres. The threephases of the mechanical response of the collagenous fibril network(Figure 3.1) in skin were simplified into two phases (Figure 4.3) tominimize the number of parameters required in the optimization.This simplification was considered as being acceptable; this wasconfirmed by the small error between the experimental tests andthe model predictions. The superficial fascia of the adipose tissuewas not implemented into the models, e.g., with radial fibres. Pre-sumably, neo-Hookean hyperelastic model was sufficient to mimicthe adipose tissue.

The neo-Hookean hyperelastic model can also limit the generalfeasibility of the method, since it is not optimal at very high strainsand lacks viscoelasticity [7]. In study I, the use of Ogden hyperelas-tic material was tested, but the difference between the responses ofOgden and neo-Hookean models was less than 1 %. Even thoughthe Ogden model is more accurate for very large strain measure-ments, it would have added one extra parameter to be optimizedbut conferring no extra value to the model. For more complex ma-terial characterization and successful optimization process, severaladditional experimental measurements would be needed.

The friction between the skin and the experimental instrumen-tation could not be measured and thus the contact was assumed tobe frictionless. However, the significance of the friction was simu-lated. In study I, using the FE model, simultaneous variations inthe indenter and the reference plate friction coefficient between 0.1and 0.9 altered the mean indenter force by 0.4 - 2.9 %. Variationsin the indenter friction altered the mean indenter force by less than0.1 %, suggesting that friction effects are minor in the indentationmeasurements. In studies II-III, compared to the situation with azero friction coefficient between the skin and the suction head, 0.1,0.3, 0.5 and 0.9 friction coefficients introduced from 1.3, 2.7, 5.4 up


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to 6.7 - 9.6 % difference in the negative pressure response, respec-tively. Therefore, the friction effects may not be negligible in theexperimental negative pressure measurements.

In studies II-III, up to 0.7 % variation in the modeled responsewas explained when the suction head was modeled to be an elas-tic (parameters for the elastomer provided by the manufacturer: E= 2.83 MPa, ν = 0.49), instead of a rigid material. This suggeststhat the assumption of a rigid suction head is an acceptable sim-plification. In study III, the significance of the contact between thesuction head and the skin was simulated with several initial com-pressive skin deformations (before suction) under the suction head(0.2, 1.2 and 2.2 mm). Compared to 0.2 mm, additional 1 and 2 mmcompression led up to 20.5 and 37.4 % lower tissue deformations,respectively. This suggests that light compression might be moreuseful when maximizing the tissue deformation and the free fluidtransportation in porous soft tissues when conducting continuoussuction treatment.

The effect of changes in the Poisson’s ratio of skin, adipose tis-sue and muscle was analyzed. Typically, changes in the Poisson’sratio (0.35 - 0.45) induced small mean error in the modeled re-sponse. Therefore, realistic variations in the Poisson’s ratio usuallyinduce only minor changes in the response of the model.

In studies I and IV, ideal perpendicular indentation measure-ments on the forearm soft tissues were simulated. In the futuremodels, non-optimal measurements with angular errors in inden-tation approach should be analyzed in order to reveal some of theuncertainties in practical measurements.

Pressure acts always normal to the surface. Therefore, the di-rection of the negative pressure, that is exposed on the "dome" ofthe skin surface within the pressurization area, varies. However, instudies II and III, the negative pressure field was simplified anddirected upwards (e.g. towards the positive Y-axis in Figure 7.5).The intensity of the negative pressure field was attenuated towardsthe edges of the pressurization area and the attenuation profile wasshaped as a portion of a sphere [26]. The impact of this simpli-


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fication is assumed as being minor, as demonstrated by the smalldifference between the experimental and model responses.

8.6 CLINICAL STUDY - PRELIMINARY RESULTS

A pilot follow-up study has been conducted to evaluate therapiesof lymphoedema in female volunteers (N=13). After breast cancersurgery, lymphoedema had been diagnosed in the upper limb ofthe subjects. Their physical characteristics were as follows (mean ±SD): age = 62 ± 9 years, height = 163 ± 6 cm, weight = 86 ± 17 kg.All subjects gave their informed consent.

The subjects were randomized into two groups: negative pres-sure therapy (N=7) and traditional manual lymphoedema therapy(N=6). Lymphoedema therapy was given during ten working daysover the course of two consecutive weeks. Each therapy sessionconsisted of lymphoedema therapy and application of a compres-sion dressing. The study protocol consisted of two research ad-missions: immediately before and after the treatment period. Themeasurements were conducted in the Department of Physical andRehabilitation Medicine of the Kuopio University Hospital (Kuo-pio, Finland). The impact of the therapies was evaluated with nineparameters; here the preliminary results of the circumferential andthe indentation measurements are presented.

The mean circumference of the healthy arm was 17.7 - 38.0 cmin the area between knuckles and armpit. The circumference ofthe swollen arm was 0.2 - 3.2 cm higher. The circumference of thehealthy arm decreased by 2.8 ± 0.9 % and 0.7 ± 3.0 % after negativepressure and manual therapy, respectively. Similarly, the circumfer-ence of the swollen arm decreased by 4.1 ± 1.0 % and 4.7 ± 1.3 %after negative pressure and manual therapy, respectively. The stiff-ness of the healthy arm incresed by 3.2 ± 17.5 % and 2.2 ± 24.4% after negative pressure and manual therapy, respectively. Simi-larly, stiffness of the swollen arm decreased by 9.2 ± 14.8 % and 1.9± 8.8 % after negative pressure and manual therapy, respectively.Both lymphoedema therapy methods induced positive changes in


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the swollen arm, as assessed from the preliminary results. Neg-ative pressure therapy appears to be a more effective method fortreatment of lymphoedema. However, the number of the subjectsshould be increased before one can make definitive conclusions.


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9 Summary and conclusions

In this thesis, FE modeling was used to study the importance ofsoft tissues and geometry of the mechanical hand-held instrumen-tation on the experimental response. This instrumentation is de-signed for diagnostics (indentation, negative pressure) and/or ther-apy (negative pressure) of soft tissue disorders. The developedfibril-reinforced 3D models were used to determine the optimalstiffness meter geometry in order to maximize sensitivity. In addi-tion, two different negative pressure treatment methods were com-pared. Developed models were able to mimic the indentation andnegative pressure responses mainly by altering the characteristicsof the skin and adipose tissue layers.

The main findings and conclusions of this thesis are summarizedas follows:

1. All the applied models indicate that the properties of theskin on the experimental response of indentation and nega-tive pressure is dominating.

2. Thickness of the tissues affects the ability of the indentationinstrumentation to detect changes in the mechanical charac-teristics of the underlying tissues.

3. The models indicate that the change in the experimental re-sponses of indentation and negative pressure during physicalstress and VO were primarily caused by the increased modu-lus of skin.

4. The porous model reveals that the continuous negative pres-sure treatment method (constant movement along treatmentdirection) induces greater interstitial fluid flow than the cyclic


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method (movement along treatment direction between eachsuction).

5. From a simulation point of view, a short and thin indenter isrecommended for measurement of mechanical characteristicsin skin. However, from a practical point of view, it mightbe challenging to use a stiffness meter with a short and thinindenter. Therefore, both points of view should be consideredand compromises made in the design of such instruments.

The results of this study suggest that the studied devices canbe useful in the diagnostics of disorders that modify the mechani-cal characteristics of soft tissues. The devised fibril-reinforced softtissue models may be further used to analyze the sensitivity of var-ious experimental instrumentation and to develop a sophisticatedand optimized device with a geometry that maximizes sensitivity.The FE models could also be used to study the long-term impact ofnegative pressure therapies.


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Publications of the University of Eastern Finland Dissertations in Forestry and Natural Sciences

Publications of the University of Eastern Finland

Dissertations in Forestry and Natural Sciences

isbn 978-952-61-1491-0

Jarkko Iivarinen

Diagnostics of Human Forearm Soft Tissues Using Indentation and SuctionMeasurementsExperimental and Modeling Analysis

Biomechanical methods for sensitive

analysis of mechanical properties

of soft tissues can be valuable in

diagnosing and monitoring the state

of pathologies, such as lymphoedema,

and aiding curative treatments. At

present, the relative contribution

of different soft tissues (skin, fat,

muscle) to mechanical response is

not well understood. This thesis

introduces new innovative modeling

methods to quantitatively analyze

soft tissues and measurement

instrumentation that may help to

diagnose and monitor alterations in

mechanical properties of soft tissues.

dissertatio

ns | 141 | Ja

rk

ko

Iivar

inen

| Diagn

ostics of Hu

man F

orearm S

oft Tissu

es Usin

g Inden

tation and S

uction M

easurem

ents

Jarkko Iivarinen

Diagnostics of Human ForearmSoft Tissues Using Indentation

and Suction MeasurementsExperimental and Modeling Analysis

Documents

Diagnostics of Human Forearm Soft · i i i i i i i i JARKKO IIVARINEN Diagnostics of human forearm soft tissues using indentation and suction measurements Experimental and modeling