\r\n\tThe aim of this book will be to describe the most common forms of dermatitis putting emphasis on the pathophysiology, clinical appearance and diagnostic of each disease. We also will aim to describe the therapeutic management and new therapeutic approaches of each condition that are currently being studied and are supposed to be used in the near future.
",isbn:null,printIsbn:"979-953-307-X-X",pdfIsbn:null,doi:null,price:0,priceEur:0,priceUsd:0,slug:null,numberOfPages:0,isOpenForSubmission:!1,hash:"278931ae110500350d8b64805c70f193",bookSignature:"Dr. Eleni Papakonstantinou",publishedDate:null,coverURL:"https://cdn.intechopen.com/books/images_new/7934.jpg",keywords:"Atopic eczema, Interleukin, Topical corticosteroids, Hand eczema, Blisters, Pruritus, Irritant contact dermatitis, Allergic contact dermatitis, Discoid eczema, Sebaceous glands, Inflammatory dermatitis, Facial rash",numberOfDownloads:null,numberOfWosCitations:0,numberOfCrossrefCitations:0,numberOfDimensionsCitations:0,numberOfTotalCitations:0,isAvailableForWebshopOrdering:!0,dateEndFirstStepPublish:"February 5th 2019",dateEndSecondStepPublish:"March 19th 2019",dateEndThirdStepPublish:"May 18th 2019",dateEndFourthStepPublish:"August 6th 2019",dateEndFifthStepPublish:"October 5th 2019",remainingDaysToSecondStep:"2 years",secondStepPassed:!0,currentStepOfPublishingProcess:5,editedByType:null,kuFlag:!1,biosketch:null,coeditorOneBiosketch:null,coeditorTwoBiosketch:null,coeditorThreeBiosketch:null,coeditorFourBiosketch:null,coeditorFiveBiosketch:null,editors:[{id:"203520",title:"Dr.",name:"Eleni",middleName:null,surname:"Papakonstantinou",slug:"eleni-papakonstantinou",fullName:"Eleni Papakonstantinou",profilePictureURL:"https://mts.intechopen.com/storage/users/203520/images/system/203520.jpg",biography:"Dr. med. Eleni Papakonstantinou is a Doctor of Medicine graduate and board certified Dermatologist-Venereologist. She studied medicine at the Aristotle University of Thessaloniki, in Greece and she continued with her dermatology specialty in Germany (2012-2017) at the University of Magdeburg and Hannover Medical School, where she completed her dissertation in 2016 with research work on atopic dermatitis in children. During this time she gained wide experience in the whole dermatological field with special focus on the diagnosis and treatment of chronic inflammatory skin diseases and also the prevention and treatment of melanocytic and non-melanocytic skin tumors. Her research interests were beside atopic dermatitis and pruritus also the pathophysiology of blistering dermatoses. In addition to lectures at german and international congresses, she has published several articles in german and international journals and her work has been awarded with various prizes (poster prize of the German Dermatological Society for the project: 'Bullous pemphigoid and comorbidities' (DDG Leipzig 2016), 'Michael Hornstein Memorial Scholarship' (EADV Athens 2016), travel grant (EAACI Vienna 2016). Since 2017, she works as a specialist dermatologist in private practice in Dortmund, in Germany. Parallel she co-administrates an international dermatologic network, Wikiderm International and she writes a dermatology public guide for patients, as she is convinced that evidence-based knowledge has to be shared not only with colleagues but also with patients.",institutionString:"Private Practice, Dermatology and Venereology",position:null,outsideEditionCount:0,totalCites:0,totalAuthoredChapters:"1",totalChapterViews:"0",totalEditedBooks:"0",institution:null}],coeditorOne:null,coeditorTwo:null,coeditorThree:null,coeditorFour:null,coeditorFive:null,topics:[{id:"16",title:"Medicine",slug:"medicine"}],chapters:null,productType:{id:"1",title:"Edited Volume",chapterContentType:"chapter",authoredCaption:"Edited by"},personalPublishingAssistant:{id:"270941",firstName:"Sandra",lastName:"Maljavac",middleName:null,title:"Ms.",imageUrl:"https://mts.intechopen.com/storage/users/270941/images/7824_n.jpg",email:"sandra.m@intechopen.com",biography:"As an Author Service Manager my responsibilities include monitoring and facilitating all publishing activities for authors and editors. From chapter submission and review, to approval and revision, copyediting and design, until final publication, I work closely with authors and editors to ensure a simple and easy publishing process. I maintain constant and effective communication with authors, editors and reviewers, which allows for a level of personal support that enables contributors to fully commit and concentrate on the chapters they are writing, editing, or reviewing. I assist authors in the preparation of their full chapter submissions and track important deadlines and ensure they are met. I help to coordinate internal processes such as linguistic review, and monitor the technical aspects of the process. As an ASM I am also involved in the acquisition of editors. Whether that be identifying an exceptional author and proposing an editorship collaboration, or contacting researchers who would like the opportunity to work with IntechOpen, I establish and help manage author and editor acquisition and contact."}},relatedBooks:[{type:"book",id:"6550",title:"Cohort Studies in Health Sciences",subtitle:null,isOpenForSubmission:!1,hash:"01df5aba4fff1a84b37a2fdafa809660",slug:"cohort-studies-in-health-sciences",bookSignature:"R. Mauricio Barría",coverURL:"https://cdn.intechopen.com/books/images_new/6550.jpg",editedByType:"Edited by",editors:[{id:"88861",title:"Dr.",name:"R. Mauricio",surname:"Barría",slug:"r.-mauricio-barria",fullName:"R. 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Venkateswarlu",coverURL:"https://cdn.intechopen.com/books/images_new/371.jpg",editedByType:"Edited by",editors:[{id:"58592",title:"Dr.",name:"Arun",surname:"Shanker",slug:"arun-shanker",fullName:"Arun Shanker"}],productType:{id:"1",chapterContentType:"chapter",authoredCaption:"Edited by"}},{type:"book",id:"878",title:"Phytochemicals",subtitle:"A Global Perspective of Their Role in Nutrition and Health",isOpenForSubmission:!1,hash:"ec77671f63975ef2d16192897deb6835",slug:"phytochemicals-a-global-perspective-of-their-role-in-nutrition-and-health",bookSignature:"Venketeshwer Rao",coverURL:"https://cdn.intechopen.com/books/images_new/878.jpg",editedByType:"Edited by",editors:[{id:"82663",title:"Dr.",name:"Venketeshwer",surname:"Rao",slug:"venketeshwer-rao",fullName:"Venketeshwer Rao"}],productType:{id:"1",chapterContentType:"chapter",authoredCaption:"Edited by"}},{type:"book",id:"4816",title:"Face Recognition",subtitle:null,isOpenForSubmission:!1,hash:"146063b5359146b7718ea86bad47c8eb",slug:"face_recognition",bookSignature:"Kresimir Delac and Mislav Grgic",coverURL:"https://cdn.intechopen.com/books/images_new/4816.jpg",editedByType:"Edited by",editors:[{id:"528",title:"Dr.",name:"Kresimir",surname:"Delac",slug:"kresimir-delac",fullName:"Kresimir Delac"}],productType:{id:"1",chapterContentType:"chapter",authoredCaption:"Edited by"}}]},chapter:{item:{type:"chapter",id:"52051",title:"The Role of Finite Element Analysis in Studying Potential Failure of Mandibular Reconstruction Methods",doi:"10.5772/64890",slug:"the-role-of-finite-element-analysis-in-studying-potential-failure-of-mandibular-reconstruction-metho",body:'\nThe lower jaw or mandible is the only load bearing, moveable bone of the skull. Defects in the lower jaw or mandible can happen as a result of trauma, infection or after a resection for pathology, which can be benign or malignant. The field of medicine or surgery has never been able to satisfactorily reconstruct a mandible after the original tissue has been gone. An unrepaired defect in the mandible, depending on the location of the defect would lead to (1) constriction of the remaining tissue around the defect leading to a malocclusion due to the pull of the muscles, (2) collapse of the arches leading to inability to eat and function and (3) difficulty swallowing or even breathing especially when lying flat due to loss of attachment of the tongue, sometimes leading to aspiration or asphyxiation.
\nThere have been many methods advocated for reconstruction of the mandible, most of which have never fully considered the biomechanical forces acting on the mandible both in the short and long term. These methods have been “tried”, most times with catastrophic results. Among the methods for mandibular reconstruction advocated are:\n
Soft tissue flap
Autogenous bone blocks
Mandibular reconstruction/bridging plate
Cancellous bone in titanium mesh
Vascularized free flap
Newer methods like the endoprosthesis or alloplastic replacement
Tissue engineered bone scaffold
The current gold standard is still the vascularized free flap, which needs a long operation time and harvest of tissue from another part of the body [1, 2].
\nThis brings the question of “what is the ideal method of mandibular reconstruction?” to the fore.
\nThe ideal method of mandibular reconstruction would:\n
Reconstruct the missing soft tissue and bone
Allow replacement of teeth
Not need a long surgery
Easy to learn without needing extensive training and skills that are hard to learn
Not need to take tissue from another part of the body
Cost effective
Not need prolonged hospitalization and recovery
Allow the patient to eat and function early or immediately
Be able to withstand the forces of biting and chewing permanently or for a long time
No such ideal method exists.
\nThe challenges to replacing a section of the mandible lie in its form and function as well as its unique location in the body. Any hardware (plates and screws) used to fix the mandible would undergo unique stresses not seen in other parts of the body. There is a non-axial load in that the long axis of the teeth is about perpendicular to the long axis of the mandible. The overlying tissue in the mouth is thin; any break in the soft tissue could lead to exposure of the hardware to the bacteria and saliva in the mouth, leading to the formation of biofilms. Ingestion of hot and cold food and liquids can cause expansion and contraction of the hardware, which is dissimilar to that of the underlying bone.
\nThe biomechanics of the mandible has not been very well studied. Mandibular biomechanics can be studied with finite element analysis, strain gauges or photoelastic models. Finite element analysis allows study of the forces throughout an entire structure but is limited to static forces. An understanding of the anatomy and biomechanics of the intact mandible is necessary prior to looking at how to conduct a finite element analysis of the mandible.
\nThe mandible is a U-shaped bone, connected at the temporomandibular joints at both ends to the skull. It consists of a corpus (body), the symphysis, which connects both right and left at the midline, the alveolar bone, which supports the teeth, the ramus with the condylar as well as the coronoid processes (Figure 1).
\nSchema showing the gross anatomy of the mandible.
The teeth are connected to the alveolar bone by periodontal ligaments, which act as a gomphosis (allowing minute movement). These ligaments are inserted into the bone on one end and into the cementum of the teeth.
\nThe muscular attachments can be divided into three groups:\n
Muscles of mastication—medial pterygoid, lateral pterygoid, masseter and temporalis; only the lateral pterygoid muscle assist in opening the mandible, the rest closes the mandible.
Suprahyoid muscles that assist in some opening of the mouth and swallowing—mylohyoid, hyoglossus, genioglossus, digastric muscles.
Muscles of facial expression that insert into the mandible—buccinators, depressor anguli oris, mentalis.
The outer surface of the mandible consists of dense cortical bone, the thickness of which varies. In certain areas, there is only cortical bone throughout. The alveolar processes and in the middle of the mandibular body and part of the ramus consist of cancellous bone (bone marrow). There is a nerve coursing through the mandible in a canal, which usually does not play a role in terms of biomechanics.
\nThe attachment of the mandible to the skull consists of the attachments of the muscles of mastication and the temporomandibular joint. The temporomandibular joint consists of two joint spaces, the superior and inferior joint spaces, surrounded by a capsule consisting of elastic collagen fibres and divided by the articular cartilage, which is a fibrocartilage. The movement of the joint consists of two distinct movements, which are Phase I (rotation) about a hinge for the first 20 mm followed by Phase 2 (translation), which is affected mainly by the action of the lateral pterygoid muscle pulling the entire condyle to the front and out of the glenoid fossa onto the part of the zygomatic process of the temporal bone. The forward movement of the condyle is limited somewhat by a protrusion called the articular eminence.
\nThe mandible functions as a Class III lever. During function, there is a zone of tension on the alveolar part of the mandible and a zone of compression on the lower border. Studies by Meyer [3, 4] showed bone deformation in the mandibular condyle region, with tensile stress along the anterior ramus as well as along the sigmoid notch area and compressive stress along the posterior ramus border. This suggested that there is a tendency of the mandible to straighten during function. This somewhat simplistic model holds true when there is bilateral and equal function and bite forces (Figure 2).
\nUpon contraction of the muscles of mastication, the mandible is bent in a sagittal plane; this is produced by the vertical component of the muscle forces, the joint reaction forces and the reaction forces from chewing motions. During asymmetrical loading (biting on one side), the largest shear forces happen between the bite force and the muscle force on the working side (the biting side) and between the muscle force and joint force on the balancing side (non-biting side). This produces then a converse load distribution with a zone of tension on the lower border and a zone of compression on the alveolar portion in the working side and vice versa on the balancing side. This means that during incisal biting (biting on the front teeth), there is an equal amount of sagittal bending on both sides but a different deformation on the working and balancing sides during molar biting [5–7].
\nZones of tension and compression in the mandible.
There is also a tendency for narrowing of the mandibular arch from parasagittal and transverse deformation upon clenching and incisal biting. This is caused by bilateral torsion of both mandibular bodies and bending at the symphyseal region leading to compression at the superior margin of the symphysis and tension at the inferior margin.
\nHylander showed that the mandibular symphysis undergoes three distinct patterns of stress and deformation, that is, corporal rotation (relative outward rotation of both halves of the mandible), medial convergence (change in mandibular width during function) and dorso-ventral shear (movement of both mandibular halves relative to one another in the vertical plane) [8] (Figure 3).
\nThere is also some difference in the deformation between the outer (buccal) surface of the mandible and the inner (lingual) surface. Lateral transverse bending occurs and the bending moment increases from back to front during the late power stroke of biting/clenching. The maximum magnitude of the bending occurs near the symphysis. This bending produces compressive stresses at the buccal cortex and tensile stress at the lingual cortex. The deformation has been calculated to be as large as 0.6 mm in a simulated molar bite of 526N using finite element analysis [6]. The mandible deformed in a helical pattern upwards and towards the working side, with regions of high tensile stress (15–25 MPa) from the coronoid process and ramus towards the lingual side of the symphysis. The highest value of compressive stress (15–25 MPa) was at the bite point and bilateral sigmoid notches, at the working side angle and in an area from the posterior surface of the balancing side ramus running to the lower border of the body till the symphysis. This then runs up to the buccal side from the inferior until the bite point. Overall, the shear stresses were larger on the working side with the exception of the balancing side condyle (peak shear stress of 25 MPa). In a nutshell, this means that the mandible changes in dimensions during function as a result of its shape and muscle pull with the greatest change in dimension and deformation in the midline.
\nWhat does all this information mean and what is the practical application? It means that application of any hardware to the mandible must take into account these forces and change in dimension. This has led to the creation of Champy’s ideal lines of internal fixation for fixation of a mandibular fracture [9] (Figure 4). Placement of bone plates and screws in the area between the zones of tension and zones of compression will tend neutralize the forces and stabilize the bony fragments enough for healing. This however uses the principle of cross bracing and load sharing, that is, using inter-fragmentary bone friction to help stabilize the bony segments.
(A) Tendency for mandible to straighten as well as undergo torsion. (B) Forces acting about the midline of the mandible. CR – corporal rotation; MC – medial convergence; DVS – dorso-ventral shear.
In comminuted fractures or in a segmental mandibular defect, the principle of load sharing cannot be applied. We are then dependent on using load bearing bone plates, which means the material strength of the plate is the only thing keeping the mandibular segments together. The method of fixation of the bone plate to the bone also plays a factor.
\nChampy’s lines for ideal internal fixation – placement of bone plates along these lines will stabilize the bone fragments enough for healing of fractures.
A very accurate finite element model is practically impossible to create due to the complex anatomy of the mandible. A true to accurate model that takes everything into account would need several supercomputers. Assumptions will have to be made to simplify the model and reduce the need for computing power.
\nSeveral authors have created finite element models of the mandible, each with different levels of complexity [10, 11]. Following are the steps taken to create an accurate finite element model [12].
\nThe information about the external shaped of the mandible is needed to be able to create a mesh to input into the FEA software. Several possible methods that have been used to get the geometry of the mandible:\n
Digital creation of 3D model—not accurate representation of the anatomical detail. The question then arises: “At what level of intricacy would there be detriment to the level of accuracy?” Is it necessary to recreate all the intricacies (concavities, canals)?
Conversion of digitized slices or sections of a human mandible into a whole 3D structure [13].
Computer tomographic (CT) scans of a human mandible or mandibular equivalent and conversion of the radiographic images in Digitised Communication in Medicine (DICOM) format into a 3D structure in Standard Tessellation Language (STL). This information is then meshed with any number of software programs. This is the most popular method. Some authors have also used the cone beam CT, which is another way to get the CT images with a cone beam instead of a fan beam. This tends to produce many sharp triangles, which need to be simplified prior to mesh creation.
Once the 3D geometry has been obtained, it is subdivided into a finite, large number of geometrically simplified elements, connected together at the nodes. The mesh is a contiguous collection of these simple-shaped elements. Most FE software have automated mesh generation features, creating relatively dense meshes, which can be refined in different regions.
\nThe material properties of the elements, namely the elastic modulus and Poisson’s ratio must be defined. Since the mandible consists of cortical and cancellous bone together with the teeth, several assumptions need to be made to simplify the model further. It is well known that bone is anisotropic in different dimensions. For purposes of simplifying the calculations, most authors have tended to assume that bone is isotropic and that the mandible is purely cortical bone. Any teeth present, which in real life would have periodontal ligaments and allows minute movements would tend to be assumed to be ankylosed to the bone, that is, fused to the bone. The teeth are composed of enamel on the outer surface of the crown followed by dentine and the pulp (a hollow cavity which contains the nerve fibres and blood supply). The outer surface of the root is covered by cementum. All these tissues have different material properties. The values for the material properties are have already been determined by studies. Some errors abound when it comes to the values of the cancellous bone. Most studies have tended to remove a block of cancellous bone and then subject the block to mechanical testing to ascertain its material properties. There is evidence to suggest that this is not entirely accurate. Misch et al. (1999) showed that the presence of cortical bone increases the elastic modulus of cancellous bone [14]. When the cortical bone was present, the elastic modulus ranged from 24.9 to 240 MPa (mean 96.2 MPa). When cancellous bone only was tested, the elastic modulus reduced dramatically (3.5–125.6 MPa). This means that the values from the literature for cancellous bone are not accurate.
\nBoundary conditions are important to prevent movement of the individual units so that the model can be loaded and deformed as a rigid structure, allowing computations to be performed. It can be divided into essential boundary conditions (displacement constraints to anchor the model and the non-essential boundary conditions or loading conditions, which are the forces to be applied to the model). Decisions will need to be made also about the insertion of the muscle forces and the force of each muscle. Some muscles like the masseter have three distinct types of fibres with different vectors. Most of these values are already in the literature.
\nWith all the information, the completed model is then solved to obtain the displacements and the resulting stress and strains. In biomechanical models, what is most often sought is information on the stresses and strains, as the force is usually known.
\nThe external forces {F} and the mechanical properties/geometry {K} are used to solve the nodal displacements {D}. With the nodal displacements known, the displacement fields are then interpolated from the nodal values using standard interpolating polynomial functions. The strain distribution is the differentiation of the displacement field yields and the stress distribution is then determined mathematically.
\nValidation can be performed by evaluating the precision and accuracy of the model. Precision, defined as how close the model’s results are to the exact solution to the biomechanical model, can be ascertained by conducting a convergence test where meshes of different refinements are created and the strains/stresses at specific locations are compared. Most reported studies have tended to use precision studies as a measure of validation as it is difficult to affix strain gauges to the human subject to test for accuracy for ethical and practical purposes.
\nThere have been very few finite element analysis conducted on reconstructed mandibles, mainly due to this being a field that is not very well understood by the people who operate in this area. In the field of orthopaedics, the biomechanics of the limbs has been well studied and there are numerous studies using finite element analysis. This section reviews the few studies that have been conducted followed by the authors’ own studies. By necessity, it is impractical to go into very much detail for each study. We will concentrate on two studies in detail at the end of this section.
\nMarginal resection of the mandible is performed for tumours that affect the alveolar process of the mandible but does not extend to the mandibular basal bone. One of the complications that can occur is fracture of the remaining portion of the mandible, usually from the corners of the resection where areas of stress concentration occur (Figure 5).
\nMarginal resection of the alveolar portion of the mandible, leaving the basal bone intact.
Wittkampf et al. [13] conducted studies on prevention of mandible fracture after a marginal resection of the mandible. The authors sectioned a human mandible, photographed the slices and digitized it. Marginal resections of different radii were placed in digitally and the areas of maximum stress concentrations in the corner of the resections were compared. It was found that an enlarge radius of resection at the corners offered the best resistance to minimize fracture after resection.
\nReconstruction plates alone are sometime used to bridge a defect to maintain the space and contours of the jaw in patients in poor health or with advanced tumours. A patient might not be fit for a long surgery or it might not be worthwhile to subject a patient to a long surgery if the prognosis is poor. Sometimes also the reconstruction plate is placed together with bone grafts, which is then subject to other considerations in terms of biomechanics (Figure 6).
\nThe most common complications in using a reconstruction plate to bridge a defect are plate fracture, sometimes after a few years, loosening of the screws and exposure of the plate (dehiscence) either in the mouth or through the skin.
\nFor pure mandibular bridging plates alone, the size and location of the defect and whether the defect crosses the midline (symphysis region) plays a large role in terms of complications [11, 15]. Recall the earlier section where it was shown that the midline is subject to many different forces of tension, compression and torsion with changes in dimensions. The masticatory loads on the plates cause vertical discrepancies that can lead to bone resorption and screw loosening. Arden et al. [16] reported that defects larger than 5 cm of bone length are associated with a high complication rate as high as 81% when plates alone are used to repair lateral defects.
\nMartola et al. [17] hypothesized that residual stresses from bending a stiff plate (sometimes repeatedly) to adapt to the jaw contour can be a main reason for plate fracture. This makes sense from a material science point of view, in that repeated bending of a metal results in work hardening of the metal, sometime with creation of small micro-cracks which may affect the mean stress in fatigue loading.
\n(A) Reconstruction plate bridging a defect in the anterior. (B) Fracture of the reconstruction plate, a common complication. Indicated by arrow.
Kimura et al. [18] investigated the most suitable method in dispersing stresses around the screws in plate fixation to the remnant mandible after a resection of the mandible. The authors took CT scans of dry human mandibles and created eight digital edentulous (no teeth) mandible models. Defects were created on the models in the front (midline) and lateral areas and the plates were drawn onto the defects with different screw configurations. A dental implant was drawn into the opposing (contralateral) side of the defect. The material properties (based on reported data) of all the components in the model were defined: cancellous bone, cortical bone and titanium. A maximum bite force of 300 N in a vertical load pattern was chosen and used. In the analysis, the stresses were concentrated around the implant, the screw closest to the defect on both sides (crucial screw) and the plate on the non-loaded side. Defects in the midline (central defects) placed greater maximum stress on the screws. If three screws were placed for lateral defects, there was greater stress on the crucial screws, which could have been due to bowing of the plate.
\nA partially dentate mandible of a cadaveric male was used to create a model with defect in the front and back. The authors (Schuller-Gotzburg et al, [19]) studied the effects and change in bone stress after caudal and buccal placement of placing a bridging plate. In this study, bridging plate alone in a conventional placement, bone grafts fixed with small miniplates to the remnant mandible and also placement of a bridging plate in a caudal and then buccal location for the reconstruction. The load was 50 N on the right mandibular second premolar. The conclusion was that there would be a better biomechanical advantage and lesser stress with the plate in a caudal position.
\nKnoll et al. [20] investigated replacement of angle defects with a standard 2.7 mm reconstruction plate with a linear screw configuration. An edentulous finite element mandible model was created; a defect placed in the right angle with a virtual bone plate placed and the model was loaded with 135 N of force in the front. The model showed that the stresses were far in excess of the material strength of titanium and cortical bone. This result showed that there was a high possibility of plate fracture, bone loss and screw loosening. The recommendation was to redesign the plate to allow screw placement in a triangular or square configuration to further maximize the interface between bone and plate.
\nVascularized free flaps are flaps of soft tissue and bone, harvested from another part of the body with its own arterial blood supply and venous drainage. The flaps are then placed into the part of the body that needs it, connected to the local blood supply and then fixed to the surrounding bone and soft tissue. For defects of the mandible, the two most common flaps used are the fibula and iliac crest free flaps. This is still the gold standard after a resection.
\nTie et al. [21] constructed finite element models to study which flaps, the fibula or iliac crest free flaps would be best biomechanically to replace a segment of the mandible. The authors scanned the mandible, iliac crest and fibula of a healthy 30-year old volunteer and used the images to create a finite element model of the mandible. Defects were created in the anterior and lateral mandible; the outlines of the defects outlined and extracted using the software and the volume of the fibula and iliac crest made to fit the defects. The stress distribution for the iliac crest was found to be similar to that of the intact mandible. The fibula reconstruction, however, had greater stresses (compressive and tensile) at the grafted bone with the maximum stress at the interface between native and grafted bone. The increased binding or interface between the iliac crest and native bone contributed towards better transmission of the bite forces. Their conclusion is that for smaller defects, the iliac crest would be better due to the above findings. As the fibula has greater length, it would be more suitable for larger defects.
\nAn endoprosthesis is an implant, mostly titanium that is placed to replace a defect in the line of the remaining bone stumps. It is usually attached to the bone stumps with a stem, which is usually cemented or press fit. This concept has been used in orthopaedics in the long bones with great success for decades [22–25].
\nThe concept of using an endoprosthesis for mandibular replacement was introduced by Tideman, initially as proof of concept and with several animal studies [26–28]. The decision was made to look into the mandibular endoprosthesis as a modular format rather than customized. The reasons for this were as follows:\n
Parts machined in large quantities are cheaper.
Customized implants require time for manufacture and are more expensive. Between the time from the scanning of the patient to the time it takes to design and manufacture as well as transport time, it may take a few weeks; this might have allowed time for the lesion if it is cancerous to grow larger in size and thus rendering the customized fit potentially inaccurate.
A stock endoprosthesis, which comes in different lengths and can be assembled in modules, allow variations and more flexibility during the surgery to adapt to defect length changes.
The modular endoprosthesis has been used, again with great success in the field of orthopaedics and musculoskeletal surgery.
\nThe difference between the long limbs and the mandible has already been discussed previously. There is also the addition of the curvature of the mandible in the anterior region, which varies between individuals, making it difficult for a stock endoprosthesis.
\nNevertheless, the animal experiments, which largely were conducted in the monkey model yielded interesting results. There were two designs: (1) the mandibular body replacement and (2) the condyle replacement. The condyle replacement had no problems with loosening and infection. The body replacement design had persistent problems with loosening between the module connections, causing infection and loosening. The cemented stems had no problems; however, a decision was made to investigate the biomechanical forces that acted on the entire reconstruction for the body replacement design [29].
\nThe design of the endoprosthesis was changed as follows [30]:\n
The stem was changed from cemented to screw retained, which was to be screwed into the marrow part of the mandible.
The module connection was changed to a male and female part in a dovetail fashion, which was connected by a screw. A slight movement of 0.1 mm was designed as s tress breaker between the connection of the male and female part.
This new design was then tested for the following aims prior to any animal or human testing:\n
Experimental evaluation of the new design to look at fatigue performance and failure patterns by mechanical testing. The entire setup was also investigated with a finite element analysis to see if the model correctly predicted the site of failure in the mechanical testing.
Finite element analysis of the new design in a simulated human mandible under certain conditions. The defect was then to be made bigger and the stem length was to be changed in length to see what happens to the stress distribution to better predict the ability of the reconstruction to withstand failure (if the stress in any part is more than the material strength of the material) and also the location of failure.
The new endoprosthesis design was made to fit the dimensions of a human mandible and underwent mechanical testing in a jig, mounted on a synthetic mandible, which had similar elastic properties to cortical bone. The methods of mechanical testing depend on the question asked. It can range from simple three-point bending test, compressive and tensile strength to complex tests for load to failure or fatigue loading. The dimensions of the assembled endoprosthesis were 18-mm stem length with 4-mm diameter and body dimensions of 15-mm length, 16-mm high and 8.5-mm thick (Figure 7).
\nThe endoprosthesis consisted of two screwed stems, which connected to each other in a dovetail. This was locked together by a central screw, which is inserted from the top.
Static testing revealed a tendency for the screw stem to pull out of the substrate of the synthetic mandible. Cyclic testing was then performed for up to 500,000 cycles. This revealed a tendency for the endoprosthesis to fail with fracture or bending at the superior surface of the stem but with no loosening of the module connection, which had plagued the earlier animal experiments.
\nThe line drawings of the endoprosthesis from the manufacturer were imported into Abaqus v6.10 (Simulia, Dassault Systemes, France). The stems were modelled as smooth cylinders to simplify calculations. Rectangular cuboids were modelled as the synthetic mandibles and bores made for stem insertion. The stems of the endoprosthesis were perfectly tied to the bores of the holes as well as the central connection screw. The cuboids were assigned the elastic properties of cortical bone and meshed with linear tetrahedral elements. It was assumed that the bone was isotropic. A bolt load of 10 N was applied to simulate tightening of the central screw. A downward force of 150 N (calculated at 80% of average static load to failure of 185 N, this was the maximum force used for the fatigue testing) was loaded on to one end of the reconstruction while the other end was given fixed boundary conditions. The load was kept constant to identify peak stresses, which lead to fatigue failure (Figure 8).
\n(A) Rendering of endoprosthesis mounted in bone blocks. (B) Von Mises stress distribution.
The finite element analysis of the setup showed areas of high stresses accumulating in the screw hole of the connection screw as well as on the superior surface of the stems. The von Mises stress recorded a maximum of 188.838 MPa, which is way below the strength of titanium alloy at 897 MPa. This corresponds well with the results of the fatigue testing with had failure of crack lines in the same areas. The conclusion of this bench top experiment was that although the forces are way below the material strength of titanium, micro-cracks as well as areas of stress concentrations from the indentation of the screw threads can lead to eventual failure over a long time of loading at a much lower load level. The connection problem of the modules seemed to have been solved.
\nA human-sized mandible synthetic was scanned with a cone beam CT scan to get the geometrical information of the mandible [31]. A synthetic mandible was used due to biohazard concerns. Due to the nature of the cone beam CT, using the DICOM information without alteration tended to produce a mesh with a lot of sharp, irregular and thin triangles. This was re-meshed with 3-Matics (Materialise, Belgium) into linear tetrahedrons. The mesh was dense enough to justify the use of linear elements to save on computational resources. The teeth were also removed digitally as it does not contribute structurally to the mandible.
\nThe endoprosthesis was modelled as previously described above (termed Case I) and then the stem length was shortened (Case II) followed by Case III where the length of the body was doubled to 30 mm (Figure 9).
\nEndoprosthesis in mandible model.
A standardized defect as well as boreholes was created to fit the dimensions of Case I, II and III on the right side of the mandible using Abaqus. The model was assumed to be made up of only cortical bone, the bone was assumed to be isotropic and the values and vectors of the muscle pull as well as the joint reaction forces were taken from the literature. The boreholes and stems were assumed to be bonded just like the previous FEM study of the experimental setup. Case I was used as the model for the standard endoprosthesis design, Case II and III were used as models for looking at the effects of a decrease in stem length as well as an increase in defect length, respectively. A 300 N load was applied directly in the incisor region. Several studies from the literature have supported our assumptions to be relatively accurate while saving on computational resources. As a form of validation, a convergence check was conducted on a mandible with quadratic elements, which gave a finer mesh. The differences were less than 10%.
\nThe analysis was conducted and divided into three separate parts:\n
Stress in the endoprosthesis
Stress in the mandible
Deflection of the mandible
Under the prescribed loading conditions, the intact mandible bent upwards almost equally on both sides. As this model was not totally symmetrical, there were minute differences on both sides. With a defect in place, reconstructed with an endoprosthesis, the mandible became less stiff, causing the left intact side to arch less than the right reconstructed side. This led to the mandible shifting to the left by 0.354 mm. There was a tendency for the endoprosthesis to bend outwards. There was little difference between Case I and II, which led to the conclusion that a case could be put forward for a shorter stem to reduce the amount of hardware needed. A longer/larger defect, as in Case III led to a tendency for separation at the module connection, leading to a possibility for eventual loosening of the connection screw in the long term. Although there was a tendency for separation at the module connection, the stresses in the connection screw were all below 100 MPA, below the material strength of titanium. The stress distribution for Case I and II was similar, but with slightly larger magnitudes in Case II. In Case III, however, since it was a larger defect, the endoprosthesis tended to bend more at the stem, although, again, the stresses were below the strength of titanium at 898 MPa (Figures 10–12).
\nWith regard to the stresses within the mandible, areas of stress concentration were in the left condyle (due to unequal deflection of the mandible from torsional stress), lower border of the mandibular body close to the abutment with the endoprosthesis and top edge of the hole in the stumps. There tended to be a pull out tendency of the stem from the holes, as experienced in the experimental setup. This was worse in Case II and III, with the peak stress exceeding the material strength of cortical bone at 85 MPa. The condyles were restricted from moving in this model, while in real life, there would certainly be movements that could dissipate stress; thus there is some mitigation factor in vivo. Whether this would be true could only be answered in animal experimental models.
\nDeflection of the mandible was measured with respect to an x-y-z axis in three dimensions at points of interest. Since the boundary conditions were applied to the incisors and condyles, this caused any deflection to show up in the body of the mandible. The greatest displacement for Case I was 0.638 mm, for Case II 0.8 mm and III 0.608 mm.
\nThe conclusions drawn from the study were that the modular endoprosthesis in its current dimensions should be adequate for small defects. Altering the stem dimensions by shortening it showed a slight increase in magnitude but no significant alteration to the stress distribution. A larger defect, however, would be more difficult to reconstruct and more studies needed to be done.
\nThis work was followed by Pinheiro and Alves [32] who performed a finite element analysis for an endoprosthesis that was not modular and comprised of a solid component, which was customized. This was a feasibility study in which the authors removed the screw stem component and designed a stem that looks to be press fit. The endoprosthesis performed well based on the finite element study. There was no tendency for separation of the module due to the entire prosthesis being a solid framework and the stress distribution as well as the displacement field were very similar to that of the intact mandible and yet did not exceed the material strength of titanium.
\nCase I Von Mises stress distribution.
Case II Von Mises stress distribution. The stem length has been halved.
Case III Von Mises stress distribution. The defect length has been doubled.
Finite element analysis is a good method to analyse solid mechanics. It has been used extensively in studying the forces in the long limbs as well as the bone plates and the pattern of bone resorption and formation. In the field of head and neck surgery, there are much fewer studies. A number of studies have been recently published as surgeons notice problems with their methods of reconstruction. We have looked at some of the studies and although not exhaustive, it should serve as an illustration to how it is used to look at potential failure of mandibular reconstruction.
\nThe objective of this review was to assess whether there is a correlation between hindlimb proximal suspensory ligament desmopathy (hindlimb PSD) and sacroiliac dysfunction (SID), and provide an understanding of the current thought process of examining these disorders. There are several studies examining the coexistence of back pain and poor performance, however for the most part, the discussion focusses on the efficacy of diagnostic techniques of the thoracolumbar region with some recognition of influencing factors [1, 2, 3]. Some authors have assumed a correlation between the two disorders in their treatment programmes [4, 5] but none quantified the association or correlation of the two conditions. There are limited studies that have looked at the structure of the sacroiliac region and applied those principles to locomotion [2] however there are many text books that describe the structure alone [6, 7]. This chapter explores the two conditions and explores the background and present theories behind hindlimb PSD and SID.
The sacroiliac joint lies deep within the pelvis of the horse, made up of the sacrum (five vertebrae fused together) and the surrounding ligaments. It is known as an atypical synovial joint [2] and a cartilaginous joint [7]. The iliac surface has fibrocartilage coverage, with the sacral surface lined with hyaline cartilage, thus creating a modified symphysis [8]. There is great variation in the joint form from L shaped to C shaped either being relatively flat or concaved, although most are at an angle of 30° [2].
The sacroiliac joint lies between the ilium wings, forming a synchondrosis that is held in place by a multitude of ligaments. These ligaments are called the dorsal and ventral sacrosciatic ligaments and the broad sacrotuberous ligament [7]. The dorsal sacrosciatic ligament has two elements, a band that runs from the dorsal tuber sacrale to the apex of the sacral spinous processes; with the lateral dorsal sacrosciatic ligament running from the tuber sacrale and ilial wing to the sacral crest on the lateral aspect. The broad sacrotuberous ligament runs from the sacrum and transverse processes of the 1st and 2nd caudal vertebrae to the ischiatic spine and tuber ischium [2, 7]. The function of this joint is to provide a relatively inelastic structure that is capable of asymmetric pelvic deformation during movement [2, 9]. The muscle structure of the back plays significant influential roles in both anatomy and biomechanics.
The movement of the horses back differs depending on the location and mediolateral swing of body mass; dorsoventral movement is seen with the greatest being middle of the back (40–47 mm per peak per stride) with a reduction cranially and caudally [10, 11, 12]. The natural movement of the lumbosacral area and the hindlimb produce a sinusoidal movement of no more than 4° within each stride cycle. Extension within this sinusoidal curve starts just moments before ground contact with the hoof, with the hindlimb at maximal protraction. In the sound horse this means that movement of the sacroiliac joint is minimal as longissimus dorsi is inactive in the impact and support phase of the flight arc of the hoof, in theory resulting in a stable joint [12, 13, 14]. Having said that linear regression revealed a significant deviation in movement over Lumbar 1 and Sacral 3 correlated to increasing speed [12]. This indicated that the movement of the back and sacroiliac joint is complex [2] and changes with every change in pace (Figure 1) [11].
Schematic of the right lateral view of the pelvis showing the position of the sacroiliac joint between the wings of the ilium and wing of the sacrum and the sacrotuberous ligament (adapted from [2]).
The movement within the joint is assumed to be little [15] due to the middle gluteal and surrounding ligaments holding it in place. Despite this, a series of studies of the human sacroiliac joint revealed adaptations to forces transmitted through the joint; which was seen as roughened areas on the contrasting surfaces [16]. Comparable studies of the equine sacrum have looked at nutational forces to determine the degree of movement and suggested there is limited movement [2]. However, another investigation raised the interesting point that when the sacrotuberous ligament was cut there was a marked increase in movement [2]. This would seem obvious, as its function is to reduce movement but does suggest that ligament damage or laxity could cause increased asymmetrical movement which in itself could have an adverse effect on the soft tissue structures of the distal limb.
The structure of the third interosseous muscle, also known as the suspensory ligament, the middle interosseous muscle or the interosseous ligament, is relatively straight forward. It originates from the proximal palmer surface of the metacarpal bones, running distally where just proximal to the sesamoid bones it bifurcates inserting on to each of the two sesamoid bones. From here it travels as the extensor branch joining the common digital extensor tendon. Even though it is termed a muscle, it is believed that once the horse matures it becomes completely collagenous in nature [7]. However, this is an over simplification as others describe the ligament as having a reduction of muscle fibres [17], while still retaining some which reduce with increased age [18, 19]. Muscle fibres quantitation showed a difference of 40% between the Thoroughbreds and Standardbreds with the Thoroughbred having less muscle fibres than its counterpart, with more muscle content being found in the hindlimb suspensory ligament than the forelimb [20]. It was also noted that the proximal region of the suspensory ligament contained less muscular tissue [19, 21]. This work also showed that the number of muscle fibres reduced with increased work intensity, thus suggesting that the suspensory ligament becomes less elastic and more susceptible to strain with increased work load (Figure 2).
Schematic left lateral view showing the interosseous ligament of the hindlimb (adapted from Budras et al. [6]).
The composition of the interosseous muscle is something of a hybrid, with the majority being collagen fibres but approximately 10% being type I muscle fibres and less than 5% type II muscle fibres. The suspensory ligament is defined by the infrequent fibroblasts embedded in the collagen matrix. These fibres are dispersed differently throughout the length of the ligament. Proximally, they are grouped as loose fascicles medially and laterally with the greater concentration just below the surface. As it reaches the three quarter mark they become less distinct, fewer in number with reduced striations. Interestingly these fibres are arranged pinnately between 45 and 80° [17, 22, 23] leading to theories that high forces are created because of the greater pinnate angle in order to stabilise the joint and indications that its purpose is anti-fatigue and postural support [24]. This was supported further by the suggestion that the elasticity of the lower limb, creating a vibration of 30–40 Hz, needs damping to reduce the likelihood of damage to tendons or bones and that this is achieved through these short muscle fibres [25, 26]. Due to the elastic nature of the suspensory ligament, it is unable to cope with sudden surges in force and is not built to deal with increased amounts of fatigue [27, 28]. It has also been noted that as the age of the horse increases so does the stiffness of a tendon unit which in turn could induce a change in kinematics [29].
It is commonly understood that tendons and ligaments play an important role in elastic strain energy during locomotion. Humans and ungulates have evolved to have more efficient locomotory systems; with equine evolution determining the distal limb muscle mass would not only be challenging to manoeuvre but very costly in terms of energy expenditure. Thus we see tendons and ligaments in the distal limb as a means of storing elastic energy [25, 28, 30, 31, 32, 33]. In order for the horse to utilise this mechanism within the suspensory ligament the energy from the ground reaction force is stored as strain energy to retract the limb [27, 32] helping to produce the break over point [34].
The function of the suspensory ligament is to stabilise the metacarpophalangeal joint and hindlimb in preventing hyper flexion in locomotion but also to act as part of the stay apparatus in preventing collapse of the fetlock joint when immobile [35, 36] effectively acting as passive control [17, 28]. However, the suspensory ligament differs slightly in its role compared to the other tendons of the distal limb. For example, the maximal stress the superficial digital flexor tendon (SDFT) and deep digital flexor tendon (DDFT) functions at is 40–50 MPa (mega-pascal units) compared to the suspensory ligament functioning at 18–25 MPa when in gallop; of course this is maximal output and decreases with decreased speeds. To gain a relative perspective, muscles work at 200–240 MPa. By comparison this seems quite small but provides an elastic energy saving of 25% for the suspensory ligament and 40% for the SDFT and DDFT which translates into an energy saving of 1.23 J/Kg at trot and 6 J/Kg at the gallop [33]; thus reducing metabolic expenditure [25, 31].
Biewener [31] calculated the peak activity stress mean standard deviation on the fore and hindlimb suspensory ligament with 53 ± 14% from walk to trot and 23 ± 19% into gallop. When ground reaction forces are considered and coupled with an increase in pace, the change in stress has an astonishingly small mean of 4%. This could be due to the kinematic calculation methods or potentially due to the biomechanical nature of the suspensory ligament. As the hoof makes contact with the ground, the suspensory ligament briefly stretches as a reaction to the ground reaction force and the sinking action of the metacarpal-phalangeal joint. The ligament then shortens to create an anti-hyperextension force. This elastic strain energy depends greatly upon the tendon shape and volume. These are varied as the suspensory ligament bifurcates distally resulting in a greatly reduced cross sectional area, leaving it under greater stress and strain [25, 31]. The elastic property of the lower limb is also heavily influenced by the individual gait pattern of each horse.
In order to understand the process of veterinary examination and its resultant observations; it is imperative to fully understand the kinematics and kinetics of the locomotion of the horse. The structure and function of the cursorial musculoskeletal systems have evolved to provide structures and patterns of movement that favour acceleration, manoeuvrability speed and endurance [30, 37, 38] which has been harnessed over centuries for various disciplines such as racing and dressage.
It is also important to note the influence central pattern generators (CPG) and proprioception have on the biomechanics of the horse. The regulated rhythm of a pace is created by the CPG neurons which are capable of generating the stimuli and therefore a rhythmic motor behaviour. Even though some believe that the CPG neurons are capable of producing this regulatory rhythm without stimulus, sensory feedback is still required [39, 40]. Minute differences the timings or intensity of these impulses of the right and left central pattern generators cause asymmetrical movement [41]. Horses that have modified their locomotory movement in an attempt to compensate for discomfort or pain of either hindlimb PSD or SID will in effect cause the CPG neurons to adapt their “pacemaker” like outputs; thus creating a new norm for the horses locomotion [38].
Locomotion occurs as a result of torque at the hip joint [42, 43] and ground reaction forces exerted on the hoof which in gallop can be as much as 2.5 times the horses body weight [44, 45], with equal magnitude working in the opposite direction providing propulsion [46]. Therefore, it is worth considering the kinematic pattern of hoof placement, to determine how the pathology of SID and hindlimb PSD may occur. The structure and function of the cursorial musculoskeletal systems have evolved to provide structures and patterns of movement that favour acceleration, manoeuvrability speed and endurance [30, 37].
The hoof does not hit the ground with a total sole impact, but instead, as a measure of control, impacts the ground with the lateral edge. This reduces the concussive effect of the initial ground contact [47, 48]. It is important to remember that the hoof at ground contact is moving forward and downward during the initial loading phase [38]. The degree of impact when the hoof hits the ground is determined by several factors; the 57:43% split of vertical impulse for fore and hindlimb respectively [23, 38], the hoof mass, size and shape of the hoof, contact surface, type of shoe i.e. racing plate or hunter with or without grips or studs. These all influence the vertical and horizontal hoof velocity, and degree of slip [37, 38, 49]. The degree of lameness also has a large influence on interplay between hoof and ground reaction force [14].
Several studies have analysed hoof velocity [38, 44, 50], two of which have considered horizontal hoof velocity of fore and hindlimbs; one demonstrating the greatest being in the non-leading limb [49] and other the leading limb [51]. The hoof velocity and leading limb has important implications to the structures in the hindlimbs; if it is the forelimb the majority of the velocity will be absorbed by the thoracic sling, if it is the hindlimb the velocity can only end at the sacroiliac joint, although this is greatly simplified. Having said that, longitudinal velocity reduces (regardless of limb) as the horse starts to break in early stance phase. In this early phase the hindlimb suspensory ligament (third interosseous muscle) is at its peak inertial capacity to prevent hyper extension, while at the same time the pitch avoidance movement of raising the head and neck backwards increases forces on the pelvic limb, as the weight is shifted backwards in the late stance phase. This increases propulsion of the moment arms of the hindlimbs, creating oscillating forces though the hindlimb [28, 52]. These oscillating forces are created with hoof-ground impact causing the limb to vibrate in a craniocaudal movement at 30–40 Hz, the greatest impact being distal in the limb. The muscles of the hindlimb act as adequate shock absorbers however risk of soft tissue damage increases with the increase in loading cycles [26]. This suggests that the greater the work load and discipline level of the horse, the more likely they are to sustain an injury. One method of removing force is slipping or sliding. The hoof is designed to allow an element of slip as a natural method of dissipating energy [53] however if sliding continues in the right conditions this can increase the risk of damage to soft tissue structures. Coupled with the ground reaction forces, this means that there are two opposing forces meeting at the horizontal axis, namely the sacroiliac joint [51].
There are many variable factors when considering the relationship between hindlimb PSD and SID; one of which is the natural biological variation in every horse, in that no two are exactly the same in conformation which ultimately enhances or impedes function. Discipline desirable traits have been documented for enhancing performance, such as the warmblood breeds for dressage, with greater hock angle reducing the incidences of injury compare to those with smaller hock angles [54, 55]. However, this was refuted in a later study of 66 warmblood horses that had the supposedly undesirable tarsal joint angle of <155.50° [56]. This was agreed with in another study examining the hock angles of 194 Warmblood horses with hindlimb PSD [57]. Hobbs et al. [54] described a selection of horses that had variations between contralateral limbs conformation and those with bone morphology variance in contralateral limbs [58]. The results of these differences may induce compensatory movements in an attempt to redistribute the weight through the stride cycle. In an attempt to counter this, and stabilise the gait, the hindlimbs may start to load in a pattern similar to a lame horse. Having said that this load distribution pattern may come from the horses’ handedness. This raises the question, if the horse is not physiologically capable of creating vertical impulsion (due to straight hocks), how and where will this affect the soft tissue structures in the hindlimb?
Asymmetries come in many forms, however each will have a marked effect on the biomechanics of the horse and more importantly the ground reaction forces; in the horses attempt to maintain equilibrium [54]. Of course, this need to maintain stability has different ground reaction forces depending on breed. Elite dressage Lusitano horses had lower vertical impulses compared to their Dutch Warmblood counterparts in collected trot with a range of 1.64 ± 0.02 N/Kg and 1.90 ± 0.08 N/Kg respectively. However this evened out with a change from collected trot to passage, with minimal difference being seen. Nevertheless, the key point in this is that the centre of mass is moved closer to the hindlimbs in the higher movements. Heim and co-authors [11] demonstrated a significant difference between Franches-Montagnes stallions (n = 27) and a general populous of horses (n = 6) in the dorsoventral movement (p < 0.02) and mediolateral movement (p < 0.01) for the spine, although to say this is a generalisation of differing anatomical parts and their role in locomotion. There is also the influence of the rider to consider here; not only as their body mass is part of the calculation but as the elite rider is capable of re-balancing even the most uneducated of horses to maintain the uphill longitudinal balance that is required of a dressage horse [59]. Dyson and colleagues [60] refuted this in their pilot study of rider weight, in that the weight of the rider had a greater significance than body mass index. Although this situation is not definitive, as there are many influencing factors in this scenario. For example, the balance of the rider and the dynamics between saddle and rider, both of which have a role in distribution of forces. In essence if the rider is displaced by an ill-fitting saddle or the rider is inexperienced the horse has to re-balance itself in order to compensate [10, 44], which in itself produces compensatory locomotion. Another interesting factor relating to distribution of forces, body movement and rider interaction was demonstrated during the heavy and very heavy rider trials, as the horse demonstrated 3/8 lameness (based on the 0–8 grade lameness scale where 0 is sound and 8 is nonweightbearing) with these heavier riders [60]. The thoracolumbar width changed with weight of rider, from 3.9% with a light rider to 2.8% with a heavy rider. Heim et al. [11] noted that there was less mediolateral movement in the vertebrae when under saddle, with a difference of approximately 10 mm in the 3rd lumbar vertebrae as compared to an 8 mm difference in the movement of the tuber sacrale. This suggested that the horses may be bracing themselves against the movement of the heavier rider. However this was an observation and not a direct conclusion. It was also suggested that the interactive surface between horse and rider, the saddle, if not fitted correctly increased the mediolateral movement of the rider, which led to their conclusion that the closer contact the rider has with the horse the more likely they are to be working in equilibrium with them [10].
The conformation of the hoof capsule and the angle of the internal structures have a role to play in suspensory ligament desmopathy and limb kinematics. A significant level of research focusses on the correlation between the navicular bone angle and force applied to the deep digital flexor tendon [44, 61]. Although the research was not directed at the hindlimb suspensory ligament; their findings still shed light on this area due to the anatomical angle of bordering structure and limb kinematics. The shape of the hoof has been reported to change the kinetics and kinematics of the distal limb. Dyson et al. [61] reported that the distal phalanx to hoof wall angle and distal phalanx to horizontal angle were smallest for deep digital flexor tendon injuries at 52.27° ± 3.29 and 50.32° ± 3.70 (mean ± SD) respectively. However, it would seem there was no direct correlation between that and the angles of the hoof wall. Research suggests that optimal hoof angles for both front and back feet should be 50–55° [62]. In addition, minimal correlation between the dorsal aspect of the distal phalanx angle and deep digital flexor tendon injury has been found and the hoof wall angle was not the same as the distal phalanx angle [61], which could account for natural variation in hoof pastern axis.
The deviation of distal phalanx angle affects the orientation of the structures above it and subsequently the metacarpophalangeal joint; which in turn has the potential to cause soft tissue injuries [63, 64]. This is because the ground reaction forces are reduced delaying break-over to latter breaking phase [64] whereas the horse should have increased loading at this point [62, 65]. This has the potential to reduce the strain on the interosseous muscle but could also inhibit the elastic strain energy needed to create its passive force.
Kane et al. [63] identified 43 race horses with ruptured suspensory ligaments with lower heel and toe angles; for example the difference between the toe heel angle control group and those with suspensory apparatus failure was 1.3° less, a relatively small number in terms of angles but quite significant over the lifetime of a horse. In real terms this means that an increase in angle of 10° increases the chance of suspensory ligament failure by 6.75 times [63].
Shoeing has been used since domestication of the horse as a means to improve performance and help maintain hoof balance. The combination of farriery techniques like rolled toes, plus different types of shoe have a significant effect on the horse’s feet and their movement [34, 45, 66]. It could be assumed that the application of the shoe would only affect the gait pattern of the horse but an 11% vertical displacement of the trunk has been observed [66], which implies a physiological effect of the structures of the back over a lifetime of a horse. Different types of shoe also have varying effects on the horse [67]. The glue on heart bar increased strain of the suspensory ligament while the racing plate alone increased strain in the superficial digital flexor tendon, interestingly when packing was added to the racing plate the increased strain was seen in the suspensory ligament. Others demonstrated an increase force of 101 N between the unshod and the steel shod foot [45, 66]. However, when looking at this in greater detail it can be seen that there is a difference in kinetics between the two states. By comparison the shod foot remains medial throughout the entire stance phase putting greater strain on the medial aspect of the limb structures. This is due to the gripping nature of the steel shoe which effectively shortens the natural slip effect of the bare foot and increases musculoskeletal forces after impact, altering the dampening effect of the suspensory ligament and preventing hoof and frog expansion on impact [34]. The stride duration also increased with the application of a shoe from (mean) 694 to 706 ms as did the stride length from 2.78 to 2.82 m; with the stride protraction and retraction decreasing after the application of shoes. This was seen as the carpal joint extending later in the swing phase and the foot being behind the movement at impact [66]. The unshod foot lands medially to then shift laterally at mid stance to then move back again medially. The application of a metal shoe removed the hoofs natural cycle of wear from the equation, which proved to be beneficial for the horse when assessing the morphology of 100 feral Brumbies [68]. Increased substrate hardness and distance travelled reduced the likelihood of hoof wall flare, however a possible negative of this is the loading of the peripheral sole in locomotion as well as the expected loading of the hoof wall [68].
There are many influencing factors when taking into consideration the relationship between horse and rider; the riders ability to control their balance, the weight of the rider and the fit of the saddle, all of these factors can have an effect on the equilibrium and the physiology of the horse. The influence of rider weight on horse movement has also been investigated. Riders were classified as light, medium, heavy and very heavy; all of which were classified as experienced riders [69]. Horses were subjectively and objectively observed with inertial sensors to determine movement at the poll and pelvis, each horse was then assessed with each rider. All heavy and very heavy rider assessments were abandoned due to temporary lameness inducement, suggesting a biomechanical change with the introduction of a dynamic load. In a study that used a lead weight added to the saddle they found the addition of weight extended the spine [70]. Thoracolumbar width changes have also been observed in another study, differing by 7.3% from the lightest to heaviest riders [71]. Variables such as saddle fit were accounted for by Master Saddlers checking prior to the tests being ridden and on the days of the test being ridden. However oscillation of the saddle in trot was reported with all rider weight groups; very heavy 14.0%, heavy 50.0%, medium 76.9% and light 84.6%, although there was no depth of discussion as to the occurrence of this except to say not all saddles fitted perfectly. Saddle bounce also occurred with the very heavy rider on 4 out of 6 horses, although this was associated with the horse being crooked in canter. Having said that, in the objective gait analysis a pelvic minimal difference of 2.2 ± 4.8 (mean ± SD) was observed [72].
Influential factors also include rider height and leg length, as this affects the fit of the saddle for both horse and rider, plus the rider’s core strength for which it is assumed that an increase in core strength would reduce rider movement in the saddle. One of the stark conclusions drawn from this study was that lameness was observed in most of the horses when being ridden regardless of rider weight (that was not apparent in hand) and that the heavier riders consistently induced severe lameness [71, 72]. This research did not answer the question of rider weight ratio but it highlighted the importance of a well-fitting saddle and the role that it plays in maintaining normal gait patterns for that horse.
An important consideration is also the discipline of the horse and the movements they are required to perfect. An example of this was elite dressage horses which are required to produce collection; “maintaining impulsion from behind to allow a lighter shoulder”, to carry out higher level movements thus distinguishing the important factor of higher proportion of bodyweight carried by the pelvic limb [73]. Although this was recognised there was no appreciation that the movement must originate in the sacroiliac joint. Furthermore the link between tarsal joint compressions was made but not associated to orthopaedic injury. However this point was contradicted by the description that the greatest movement of the SIJ to be on the transverse plane [2]. This allowed for a wider overall viewpoint comparing the likelihood of SID by disciplines; with dressage horses and show jumpers being more susceptible [2]. This suggested that SID is induced by the greater degree of collection required of each discipline and increased angles of the moment arms of the hindlimbs, in effect reducing stability of the joint.
Data analysis primarily segregates elite and non-elite horses in order to classify gross morphology [73], demonstrating the understanding that each discipline has a differing physiological impact. This is then subdivided to location or type of injury. Conversely, they did not make the distinction in forelimb and hindlimb suspensory ligament injuries, and although there were a significant number of classifications observed, it was not stated whether these were distinct individual injuries or if the horses had sustained more than one [73]. However Barstow and Dyson [1] went a step further and subdivided their cohort into sacroiliac pain only and sacroiliac pain with hindlimb lameness; thus starting to demonstrate a correlation between the two. In comparison, others recognised the presence of other abnormalities but mainly focussed on osseous changes [74]. Dyson [61] considered an alternative perspective of tarsal conformation predisposing horses to PSD and acknowledged biomechanics as a possible influencing factor but again with no correlation to SID.
The surface that horses work on have to be taken into consideration as they directly influence the impact on hoof loading (hoof sliding and the declarative longitudinal forces) and therefore the reaction of the limb structures [38]. Surfaces vary based on their composition, a ménage situation will have a hard under layer with surface applied to a specific depth, while some race tracks will run on turf. The most important element here is the cushion depth as this has the potential to absorb some of the concussion [75, 76]. Having said that, a softer surface encourages the toe to pivot causing a rotational force on the distal limb structures [38]. In a human based assessment it was found that peak forces reduced with an increase in compliant surfaces [76]. The compliance of track surfaces has also been examined, each type of surface had a distinct effect on the hoof velocity and swing phase, with the greatest deformation coming from the most compliant surface [75]. Even though it was noted that this surface caused significant increases in stance time and angle of hoof on landing, they did not draw any conclusions from this or discuss the soft tissue implications for the horse. However, it does imply that the suspensory ligament would have to sustain its force for a prolonged period and thus potentially fatigue if longer stance time occurred. This concept was looked at in greater detail with the use of a dynamometric shoe applied to three race horses which showed that turf surfaces had a greater ground reaction force (42.9 ± 3.8 g; mean ± SEM) compared to synthetic surfaces which reduced the ground reaction forces significantly (28.5 ± 2.9 g; mean ± SEM) [77]. This implies that there will be less impact on the soft tissue structures of the hindlimb and subsequently the sacroiliac joint.
In order to gain a full understanding of the relationship between hindlimb PSD and SID, the way in which the horse works, its discipline and level, plus the rider influence and ability must be considered [73, 78, 82]. Barstow and Dyson [1] used rider colloquialisms to aid quantification of lameness; this is very subjective even when well versed in this terminology [12]. This highlights the need to be objective and specific in pinpointing lameness. Similarly another study used anecdotal evidence to support their hypothesis of sports performance level and orthopaedic injury diagnosis, suggesting that this is frequently seen in practice but not yet documented [73]. Having said that, some studies [4, 5] have noted that some horses may suffer concurrent injuries of the sacroiliac joint or proximal suspensory (respectively) but did not draw conclusions from this regarding cause and effect or relationship.
As already stated, it is difficult, if not impossible to ascertain where the pain is coming from within the sacroiliac joint; one of the possibilities is the articular surface. As the horse ages there is an increased likelihood of cartilaginous deterioration irrespective of breed type or discipline. This deterioration and possible changes may be the result of long term laxity of the surrounding ligaments [83] which in itself could cause instability of the sacroiliac joint or degenerative suspensory desmitis which would alter the gait pattern of the horse permanently [84]. Another factor, of course, could be the ground reaction forces and the impact of hard work on hard ground for sustained periods.
It is recognised that lameness of the hindlimb creates compensatory movements within the lumbosacral region [74, 85]. Signs of subtle discomfort or pain are not so easily detected. A reduction in equine motivation to work or refusing jumps or bolting with their rider can be seen [4]. However, use of inertial measurement units can make the process of assessing asymmetry objective. The assessment of 60 horses used for polo showed 36 horses (60%) demonstrated an asymmetrical movement in the head, pelvic or both [86]. Statistical analysis linear regression revealed none of these measures had a slope greater in difference than zero. This tells us two things; that inertial measures are able to quantify small asymmetries in the horse but the value of this in a lameness evaluation must be left with the veterinary professionals to interpret. In reality this technology is not commonly used in practice and the standardised approach is to use diagnostic nerve blocks to determine the area of pain. However, this is not straight forward as they need to be used in conjunction with clinical examination and imaging modalities. In fact Pilsworth and Dyson [87] described clinically sound horses receiving a palmer nerve block to have a change in gait. This was echoed by Denoix and co-authors [88] when describing the pitfalls of sacroiliac nerve blocks, in that potential error could cause a false positive. In contrast others focussed on the biomechanics of the entire vertebral column [11, 82] but limited the discussion of the limbs to kinematics. This was echoed following assessment of the dynamic asymmetry of polo ponies, which again reverberated the question of correlation and cause [89].
The need to be more specific was demonstrated by Murray et al. [73] in their results making reference to thoracolumbar and pelvis but not specifically the SIJ. Goff and co-authors [90] advanced this to identify degenerative changes of the SIJ causing poor performance. However there is no correlation to unilateral or bilateral distal limb lameness. To emphasise the need to be unambiguous Murray et al. [73] used a large sample size (1069 horses), which potentially could be representative of the equine population. However, as the study was conducted at a referral hospital it would not represent primary veterinarians seeing acute injuries or stages of disease; emphasising the need for a retrospective study of primary veterinary practices.
In a study by Barstow and Dyson [1] 296 horses were assessed for SIJ pain, of which 203 (80%) showed hindlimb lameness with 181 specifically identified with proximal suspensory desmitis (89% [94% bilateral, 6% unilateral]). Although this represents relatively small numbers by comparison to sports performance studies [73] its findings are significant and showed a direct correlation. Furthermore, the work up of the horses was carried out by the same veterinarian reducing the likelihood of subjectivity in gait analysis.
In a similar study the prevalence of orthopaedic injuries was examined, classifying the horse by injury alone [91]. Having said that, discipline was acknowledged but no relationship established; although the kinematics of the show jumper’s pelvic limb were noted. A limitation of this study was that the information was extracted from yard records rather than from veterinarian’s records. Furthermore the initial assessments were made by several veterinarians potentially providing greater diversity in objectivity of lameness detection. In contrast, a unique perspective examining the likelihood of heritable degenerative suspensory ligament desmitis in the Peruvian Paso was published [92]. Dyson [61] demonstrated an understanding of this but also questioned conformation as a predisposing factor.
All of this begs the question as to how a horse with sacroiliac dysfunction and hindlimb PSD can be identified? Generalised pain detection using facial expressions has been used for many years with infants. Langford et al. [93] took this principle and adapted it to form the mouse grimace scale for those used in biomedical research, this was hailed as a great success as a pain indicator. Miller et al. [94] developed this further to include pain behaviours. The assessment of pain has always been subjective and relative to the experience of the practitioner, formalising a grimace scale for horses [95] has made this an objective process for the equine veterinarian. There are general indicators of pain as seen in the horse grimace scale whereby an assessment of the horses facial postures are calculated on an ethogram to determine general level of pain. For example, a horse with tension above the eye alone may not be indicative of pain, but coupled with ears stiffly backwards and prominent chewing muscles, it may indicate a level of pain [95]. The facial grimace scale alone has been identified as limiting an ethogram for equine pain behaviours both ridden and in hand has been developed [60]. Importantly this study ensured its efficacy by refining its use with a “within observer repeatability study” to confirm this as a suitable tool for quantifying pain behaviours. This concept was taken a step forward in order to develop a scale for the ridden horse, for example the horse moving on three tracks in trot or canter could be an indicator of sacroiliac pain [69, 70, 71, 72, 73, 74, 75, 76, 77, 78, 79, 80, 81, 82, 83, 84, 85, 86, 87, 88, 89, 90, 91, 92, 93, 94, 95, 96]. Some other indicators are a direct reflexion of the location of pain such as bucking going into canter demonstrating pain in the sacroiliac region; however, a horse at the very start of its education may resist the rider and buck out of frustration. Having said that, persistent displays of these behaviours are a direct indicator of pain [69]. There are many more subtle signs including asymmetry of the tuber coxae and the tuber ischii that can be visually assessed by the practitioner, asymmetrical muscle mass of the superficial gluteal and holding the tail to one side can also be seen as pain indicators [97]. Saddle slip has also been identified as an indicator of hindlimb lameness with a direct correlation between bilateral and unilateral lameness (p = 0.344 and p = 0.286 respectively) [98]. This advancement could improve criteria in determining the subtle variations in lameness between sacroiliac dysfunction and hindlimb PSD.
Research in the last 10 years has focussed on poor performance and diagnostic techniques, back pain and biomechanics or suspensory ligament disease. The correlation of information to demonstrate that lameness may be from one or more sites in the horse is limited. This indicates the necessity for further studies to determine whether there are correlations between hindlimb proximal suspensory desmopathy and sacroiliac disease. Understanding whether correlations are present between the two disorders could have an impact on evaluation and diagnosis, treatment and recovery, prognostics and welfare.
Albert A. Rizvanov (
The authors declare no conflicts of interest.
IntechOpen publishes different types of publications
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