Anatomic variability of the human femur and its implications for the use of artificial bones in biomechanical testing

Introduction

Osteosynthesis, or fracture fixation, consists of joining and stabilizing the ends of a fragmented bone with mechanical devices such as metal plates, pins, rods, nails, wires, or screws. The primary idea is to immobilize the fragments of a bone fracture in order to enable bone growth and healing [1]. The most crucial need for successful treatment and thus healing of bone fractures is adequate mechanical stability. Interfragmentary shear forces are known to be deleterious to bone healing, whereas axial forces are beneficial [2].

Although modern fracture implants demonstrate greater mechanical stability in both experimental and computational assessments, this may not always translate into equal gains in fracture healing problems. For example, treatment outcomes of fractures caused by bone loss (osteoporosis) is still not satisfactory and is associated with an alarmingly high morbidity and mortality [3].

The biomechanical evaluation of new clinical implants is currently only poorly standardized. Furthermore, the comparability of biomechanical evaluations of bone-implant constructs is hampered by a number of influencing factors, such as type of loading (static or dynamic), direction of load application (generic or physiological, e.g., sit-stand-transfer or gait cycle), the implant (different companies, models, sizes, and designs), the inter-operator variability (individuality of the surgeon or placement of the implant on the bone by use of templates), the type and fabrication of the fracture (osteotomy or realistic fracture) and last but not least the variety of specimens (human fresh frozen or embalmed, artificial, healthy or osteoporotic) [4].

As bone is a heterogeneous and anisotropic material, the properties derived from mechanical testing vary depending on the type of bone (long bone, flat bone), the location within the bone (lateral-medial, anterior-posterior), or the orientation of the test with respect to the bone (along or against the main direction of the trabecular structure) [5]. Thus, the substrate on which the new implants are evaluated is one of the most essential elements in planning a biomechanical study for the development and evaluation of implants. The substrate’s major goal is to imitate the “in vivo” conditions of human bone as accurately as possible. In the biomechanical testing of osteosyntheses, either human specimens or artificial bones are utilized; both have advantages and downsides [6].

There exists significant disparity between the determination of fracture stability due to implant-treatment in experimental methods and their ability to predict the outcome of stability and fracture healing in a patient [7]. One possible explanation for this disparity is a failure to account for population-specific characteristics. Gender- [8, 9], age- [8, 10], and ethnicity- [11, 12] related differences in bone morphology and quality all have an impact on the mechanical behavior of bone [5].

Thus, the aim of this review is to address the questions whether currently commercially available artificial bones sufficiently represent the human anatomic population variability for the mechanical testing of osteosyntheses using the femur as an example.

Testing for approval and performance evaluation of osteosyntheses

The EU Regulation 2017/745 on medical devices, known as the Medical Device Regulation (MDR), came into force on May 25, 2017 and is implemented nationally by the corresponding Medical Devices Implementation Acts. The MDR defines medical devices as products with a medical purpose that have a physical effect on the human body and are used for the detection, prevention, monitoring, treatment or alleviation of diseases. Accordingly, osteosyntheses fall into risk class III as invasive, prolonged-use and non-active medical devices. Medical devices must undergo a conformity assessment procedure and meet the legal requirements before they can be placed on the market with CE marking. The conformity assessment includes a risk analysis, proof of safety and performance parameters [13].

When evaluating implants, various tests must be carried out in accordance with legal requirements and standards. For osteosyntheses, preclinical evaluations based on scientific literature, registry data and biomechanical tests that simulate the conditions of use in the body are required. Recognized test standards (ASTM or ISO) are used to characterize osteosyntheses and simulate the conditions of use in the human body and include static and dynamic load tests, but not to predict performance in the human body. In addition, physiological tests of osteosyntheses must be carried out in combination with realistic test specimens, surrogates or human bones in order to investigate the biomechanical properties and behavior of the implants under realistic conditions. This is important for several reasons:

A crucial aspect is the force transmission between the implant and bone, as well as the investigation of loosening or fracture due to localized bone overload under static or dynamic loading [5]. The problematic issue here is the unrealistic properties of synthetic bones, which do not correspond to real bones. This discrepancy leads to fractures occurring at incorrect locations or not occurring at all. Existing publications demonstrate that the failure patterns differ between real bones and synthetic bones, especially when implants are tested not in isolation but in combination with bone or synthetic bone [14, 15]. Bones and implants must be able to withstand the complex biomechanical loads in the human body. Tests with bone simulate these loads more realistically than pure material tests and thus contribute to the optimization and safety of osteosynthesis implants [16].

In summary, tests with surrogates allow a more precise evaluation of implant behavior under real conditions and thus contribute to the optimization and safety of osteosynthesis implants. The exact testing requirements depend on the risk class and intended purpose of the implant. A notified body must carry out the conformity assessment for higher risk classes [13].

Human and synthetic bone in biomechanical testing

The fundamental advantage of human specimen used as surrogate for biomechanical testing is that they represent the structural, anatomical, morphological, and mechanical features of the human body’s “in vivo” environment and, in theory, have the required population variability of these properties. The principal drawbacks of human samples are the high acquisition costs, the scarcity of donor bones, and the extensive logistical requirements (including shipping and (frozen) storage) [6]. In most cases, elderly people donate their bodies for research, so that age-related changes, such as osteoporosis, can impact the donor bones. Samples from younger or even juvenile donors are lacking [17]. The use of fresh frozen human specimens (i.e., specimens frozen to at least −20 °C soon after the donor’s death and thawed before to biomechanical testing) poses an infection risk to laboratory staff. To reduce the danger of infection, irradiation, gas sterilization, or chemical sterilization can be utilized; nevertheless, all these techniques can considerably affect the structural and mechanical properties of human bone [6, 18]. Furthermore, physical properties of human donor bones vary greatly between samples. As a result, a relatively large number of samples is required for an effective statistical evaluation [5].

Synthetic bone models, together with human cadaver bone, are regarded as the gold standard for orthopedic biomechanical testing [6]. Synthetic bones are, however, commonly utilized in surgical training [19]. The ASTM (American Society for Materials and Testing) Standard F1839-08(2021) “Standard Specification for Rigid Polyurethane Foam for Use as Standard Material for Testing Orthopaedic Devices and Instruments” defines polyurethane (PUR) as the standard material for testing osteosyntheses, specifying the composition, physical and mechanical criteria, and test procedures for PUR. Consequently, most biomechanical models on the market are based on this standard [20].

Synthetic bones can resolve the problems of human specimens noted above: they are available in vast quantities, are less expensive than human specimens, show reduced variability and pose no risk of infection. Furthermore, storage is simpler and less expensive. Unfortunately, commercially available synthetic bones have significant drawbacks: they are usually only available in a very limited number of sizes or from one side of the body, polyurethane foams are mostly homogeneous and do not have a preferred direction such as human bone, and their mechanical properties most likely only partially reflect the properties of human bone [6].

The majority of artificial bones currently in use for evaluating osteosyntheses are manufactured by Pacific Research Laboratories [21], Synbone [22, 23] and BoneSim [24]. Most frequently used for biomechanical studies are Sawbones models [25], therefore these are used as an example in this review. The cancellous portions of the Sawbones are made of polyurethane foam. The foam is enclosed in a glass-fiber-reinforced epoxy, which is designed to simulate the cortex of a bone [21]. Sawbones artificial femora have been validated against human specimens in simple tests such as axial loading, four-point bending, and torsion of the entire bones, but only against a very small number of human specimens. However, only the mechanical properties of the shaft region of the artificial bones were considered for validation [26, 27]. Sawbones however cannot accurately reflect local human bone characteristics, particularly at implant-bone interfaces [6]. Several studies have shown that Sawbones do not resemble, or only minimally mimic, the mechanical qualities of implant-bone constructs [28], [29], [30], [31].

Femur’s population diversity and representation in biomechanical substrates

Sawbones femora for biomechanical testing are available in two common sizes: “medium” (SKU:3403), and “large” (SKU:3406) [21, 25]. The “large“ model’s anatomical parameters match those of a 60 year old male with a height of 183 cm and a body weight of 91 kg with good bone quality [32], while the “medium“ model is based on a 170 cm tall, 84 kg male of unknown age [25]. Recently, there has also been an “asian” (SKU: 3447) model introduced, which is claimed to better reflect the geometry of the general Asian population [21]. Furthermore, a “small” (SKU: 3414) model is available. The cancellous bone is available in densities ranging from 5 to 20 PCF (pounds per cubic foot; 100 PCF corresponds to 1.6 g/cm3) [21] representing varying bone qualities. Table 1 summarizes the dimensions of the four Sawbones models as listed on the manufacturer’s home page.

The anatomical characteristics of femur length, neck-shaft angle, femoral anteversion, and distal condylar width will be highlighted below to investigate possible anatomical discrepancies between commercially available artificial bones and human populations.

Variability of femur length

Moosa et al. defined femur length as the distance between the highest point of the femoral head and the infracondylar plane [33], whereas Soodmand et al. defined femur length as the distance between the most distal point in the transverse plane and a parallel plane containing the most proximal point of the femur [34].

A person’s thigh length is directly proportionate to their body height. The taller the individual, the longer the thigh, and vice versa. Human heights range from 55 cm to 251 cm, according to the Guinnes Book of Records. These are the body heights of the world’s lowest and tallest individuals [35].

Body size, and hence femoral length, differs between sexes, age groups, and ethnic groupings. A woman’s average height worldwide is 160 cm, whereas a man’s average height is 171 cm [36]. Globally, Europe has the tallest men and women, with an average height of 180 cm for men and 167 cm for women, followed by Australia. South and Southeast Asians are the shortest, with women standing at 153 cm and males standing at 165 cm on average (see Table 2, [37]).

Table 2:

Average height of women and men (adapted from [37]).

Ethnicity	Men	Women
North American	177 cm	164 cm
South American	171 cm	158 cm
Central American	168 cm	155 cm
European	180 cm	167 cm
West, East and Central Asian	171 cm	159 cm
South and South-East Asian	164 cm	153 cm
Australian	179 cm	165 cm
African	168 cm	158 cm

Body height, as previously stated, is proportional to femur length. Femur length disparities exist due to variances in average body height between the sexes. In their sexually dimorphic examination of 200 specimens, Moosa et al. not surprisingly report that males have longer femurs (average 437 mm) than females (average 402 mm) [33].

Figure 1 depicts the femur lengths of the four Sawbones models (red bars) as well as findings from studies on the femur lengths of different ethnicities or genders (blue bars). The femur lengths of the four Sawbones femur models represent well the average femur lengths of different ethnicities and genders. The “large” Sawbones model is longer than all femur lengths observed in researched publications. However, extremes such as the world’s tallest and shortest men are not included.

Figure 1:

Variability of femur length. (A) Model demonstrating the length of a femur. (B) Variability of femur length (references from left to right [21], [21], [21], [21], [38], [38], [39], [40], [39], [39], [39], [39], [40], [41]. Bars and whiskers indicate mean values and standard deviations).

Variability of the center-collum-diaphyseal angle

The center-collum-diaphyseal angle (CCD) is defined as the angle between the shaft and the femoral neck. The CCD angle can range from 106 to 151° [41], with women having a much larger CCD angle than men [42]. Furthermore, Asian ethnic groups have a larger CCD angle than Caucasian ethnic groups [43]. Age is another factor that influences the CCD angle of the femur. According to Hofmann et al., the CCD angle decreases with age [44].

The CCD angle (see Figure 2) of the “large” Sawbones model, which has a CCD angle of 120°, is less than the CCD angles of all studies examined. In contrast, the CCD angle of the “medium” model, 135°, is much greater than most of those found in the populations studied. In the CCD angle, only the Asian model matches the Asian ethnicities. However, the Sawbones models cannot match the 106–151° range described by Shivashankarappa et al. [41].

Figure 2:

Variability of the center-collum-diaphyseal angle. (A) Model demonstrating the CCD angle. (B) Variability of neck shaft angle (references from left to right [21], [21], [21], [21], [45], [45], [38], [38], [46], [47], [47], [46], [39], [39], [39], [39], [41]. Bars and whiskers indicate mean values and standard deviations).

The CCD angle is a critical parameter in the biomechanics of the femur and fracture-treating osteosyntheses, as this anatomical region is vital for load transmission between the femoral shaft and head [48]. The CCD angle affects the loading of the subtrochanteric region of the femur; smaller CCD angles are associated with higher mechanical load in this region, leading to prolonged fracture healing times [49].

The CCD angle also influences the type of femoral fractures. Higher CCD angles are typically associated with intracapsular femoral fractures, whereas lower CCD angles are linked to extracapsular femoral fractures. This variation impacts the selection of osteosynthesis methods for treatment [50]. In hip arthroplasty for fracture treatment, the implant size and the CCD angle primarily influence the calcar region. Smaller CCD angles result in greater stress in the calcar area, while other femoral regions are minimally affected [50].

Variability of the femoral anteversion

The angle formed by the femoral neck and the shaft is referred to as anteversion of the femoral neck. This angle reflects the degree of femur “twisting“ and is interesting from a biomechanical standpoint, as alterations in the anteversion angle affect the joint stress in the hip due to the associated change in the length of the lever arm [51].

There are differences in anteversion of the thigh between the sexes as well as across ethnic groups. According to Bah et al., the anteversion angle in British women is approximately 13 ± 6° and in men is approximately 6 ± 8°) [38]. Schmutz et al. found that Asian ethnic groups have much bigger anteversion angles (e.g., Japanese 22 ± 12°) than Caucasian ethnic groupings (15 ± 9°) [46]. Additionally, Dimitriou et al. discovered substantial disparities in femoral anteversion between the subject’s right and left legs [52].

With anteversion angles of 8, 11, and 20°, the Sawbones models accurately capture population diversity (Figure 3). However, the “large“ model intended to model a male caucasian femur with 8° anteversion does not reproduce the male Caucasian group’s average anteversion angle of 15 ± 9° stated by Schmutz et al. [46].

Figure 3:

Variability of the femoral anteversion. (A) Model demonstrating the femoral anteversion angle. (B) Variability of femoral anteversion (references from left to right: [21], [21], [21], [45], [45], [38], [38], [46], [46], [46]. Bars and whiskers indicate mean values and standard deviations).

The femoral anteversion angle is a crucial morphological parameter for orthopedic surgeons to consider during fracture fixation and joint reconstruction procedures involving the proximal femur, as it greatly affects the biomechanical performance and outcomes of these surgeries. For instance, the anteversion angle affects the rotational stability and alignment of intramedullary nails employed in femoral shaft fracture fixation. Deviations in the anteversion angle can arise both during the locking of the nail and from rotational forces post-fixation, potentially resulting in malalignment [53]. Abnormal anteversion angles are linked to a higher risk of hip osteoarthritis, femoroacetabular impingement, and other hip issues. Proper restoration of the anteversion angle is essential for optimal biomechanics following osteosynthesis procedures around the hip. Morphological parameters, such as the anteversion angle, influence hip joint biomechanics and range of motion. Considering the patient’s native anteversion is vital for optimal implant positioning and function in joint replacement surgeries [54].

Variability of the distal condylar width

The femur’s distal end widens to generate two condyles, which form the knee joint with the tibia. The distal femur is ethnically dimorphic. Cavaignac et al. discovered that Asian ethnic groups have a much smaller distal condyle width in comparison to European ethnic groups (Asian: 76 ± 5 mm, European: 81 ± 7 mm). They also discovered gender variations in condylar width (Asian female: 72 ± 3 mm, Asian male: 80 ± 4 mm, Caucasian female: 76 ± 4 mm, Caucasian male: 85 ± 5 mm) [55]. Yan et al. found considerably smaller condylar widths in Asian ethnic groupings, with females having 59 ± 3 mm and males having 66 ± 3 mm [56].

The condylar width of the Sawbones model in size “large” is substantially greater than the other measurements published, measuring 93 mm (Figure 4). Furthermore, the Asian model of the Sawbones femur appears to be too wide (condylar width 70 mm) to represent Asian women’s distal condylar width (59 ± 3 mm [56]).

Figure 4:

Variability of the distal condylar width. (A) Model demonstrating the distal condylar width. (B) Variability of distal condylar width (references from left to right [21], [21], [21], [21], [55], [55], [39], [55], [55], [56], [56], [39], [39], [39], [39], [57], [57]. Bars and whiskers indicate mean values and standard deviations).

The condylar width of the femur significantly impacts the biomechanical outcomes of osteosyntheses in the femoral region. Variations in condyle size and shape can affect joint loading and fracture patterns involving the condyles [58]. Studies emphasize that in minimally invasive fixation techniques for comminuted femoral condyle fractures, restoring the joint surface and achieving precise fracture alignment are crucial, highlighting the importance of condylar geometry [59]. Furthermore, condylar width can affect the stability and alignment of the fracture post-reduction, thereby impacting overall fracture outcomes. Wider condyles may provide better support and stability, leading to improved outcomes, whereas narrower condyles may result in less stable fixation and potentially poorer outcomes [60].

Moreover, femoral condyle width is critical in the selection, fit, and alignment of implants in knee arthroplasty, potentially influencing procedural outcomes. The width of the femoral condyle correlates with the sizing of femoral implants, and variations in condylar width can lead to component mismatches between knee dimensions and available knee arthroplasty systems [61]. For instance, a study by Hitt et al. found significant variation in femoral fit with available standard symmetric implants, resulting in average overhang, particularly in women [62].

Summary and outlook

The purpose of this review was to determine whether the artificial bones of the market leader for biomechanical testing of osteosyntheses accurately reflect the anatomical heterogeneity of the human population.

Macroscopically, bone has two structural features: cancellous bone and cortical bone. Cortical bone shape and strength are crucial to skeleton integrity, especially in long bones where cortical bone is predominantly responsible for axial stress transfer. The microstructure and geometry of the cortical bone affect a bone’s strength and, as a result, its resistance to and frequency of fractures [10]. Long bone geometrical and architectural features vary greatly from location to location, differ between sexes, ethnicities and also depend on age [73]. As already stated above, size, geometry and material properties have an impact on the mechanical properties of bone and may be more important to the overall mechanical performance of an osteosynthesis bone construct than the technical qualities of the osteosynthesis implant itself [5].

The Sawbones model of size “large” and “medium“ correspond to the geometries of males [25, 32]. When the individual geometric parameters are compared, the length of the femur “large” is more than the reported average femur lengths of different ethnicities or genders, as is the breadth of the distal condyles. In contrast, the CCD angle is substantially less in female Caucasians than in male Caucasians. As a result, not all of the parameters gathered are consistent with a specific population.

For the “medium” model, a similar image appears. The femur length is comparable to the average female Caucasian femur, however it appears to be overly long for Asian ethnicities. The CCD angle of the “medium“ model, on the other hand, is chosen to be quite large, larger than any data reported across all ethnicities and genders. However, the width of the condyles fits the female variability studied. Again, not all geometrical factors were taken into account to fit a specific population. When the geometry of the Sawbones femur “medium” is compared, it most nearly represents the average female, Caucasian population.

In all parameters evaluated, the Asian Sawbones model conforms to the average values of Asian ethnic groups for which data was available. Unfortunately, geometry for African ethnicities is not accessible, despite research showing that African ethnicities have higher bone densities and thicker cortices [69], which affects the mechanical stability of osteosyntheses [5].

In addition, a “small” artificial femur was created, whose modest proportions suggest that it belongs to the anatomy of a child or adolescent. From a mechanical perspective, however, the small model also includes the well-known cortical structure composed of two densities of spongy polyurethane foam and glass fiber-reinforced epoxy. Because infantile or juvenile bone is softer, more flexible, and less brittle, it is presumably not represented in the mechanical properties of bone that can also exhibit so-called greenstick fractures [74]. Still, this indicates that the industry has acknowledged the necessity for more bone models.

Sawbones femora cancellous bone is represented by polyurethane foams of varying densities and cellularity. Because they are homogeneous and do not have a preferred direction, they do not exactly correspond to the mechanics and morphometry of human bone but are able to simulate human bone mechanical characteristics in simple load scenarios and tests. Numerous investigations have demonstrated that Sawbones artificial bones cannot accurately represent the mechanics and interaction of implant and bone (e.g. [15, 30, 31]). In this review, mechanical parameters such as pull-out strength of screws or fatigue strength of synthetic bones were not investigated but research was focused on the anatomical properties of femora.

Additionally, the comparison of the mechanical properties of artificial femora to human bones is hampered by the very small number of human specimens, even more so because no demographic data are known from the human comparison specimens [27, 32, 75]. Comparability is further hampered by a wide variety of mechanical measurement methods [26, 32, 76].

Non-fitting osteosynthesis can cause complications such as inadequate fracture reduction, incorrect plate positioning, the need for intraoperative adjustments, and patient discomfort, potentially leading to re-fracture. These complications compromise both the mechanical and biological aspects of bone healing. Improperly fitting implants may not provide sufficient stability, resulting in mal- or non-union of fractures and necessitating revision surgery, thereby increasing complications and costs [77, 78].

Intraoperative adjustments, such as bending, are often required to fit a non-conforming plate, prolonging surgery and increasing risks. Poorly fitting plates can cause patient irritation and discomfort, often leading to plate removal and additional surgery. Non-fitting intramedullary nails may migrate, causing pain and discomfort [77, 79].

Non-fitting plates can fail due to bending or breaking, often linked to bone malalignment or delayed/non-union. This failure occurs if the implant endures unintended loads or if its structure is compromised by intraoperative adjustments [77].

Exemplarily, Sarai et al. report in their study on the influence of ethnicity on proximal femoral nail protrusion that the femoral length of Asian ethnic groups is significantly shorter than that of Caucasian ethnic groups, resulting in an increased incidence of intramedullary nail protrusion in Asian ethnic groups [43]. Hwang et al. investigated the discrepancy between the application of an anatomically preformed femoral plate and 60 Asian human specimens and showed that the plate protruded up to 32 mm from the femoral shaft. There was a consistent pattern of malalignment across all specimens and the authors describe that an attempt to screw the plate to the bone intraoperatively can lead to valgus malalignment at the fracture site [80].

Overall, these issues with non-fitting osteosyntheses lead to increased costs from revision surgeries, longer hospital stays, and pseudoarthrosis [77]. Mismatched osteosyntheses cause intraoperative complications, extending operative time and associated expenses highlighting the importance of considering body or anatomical dimensions when selecting osteosyntheses.

Limitations

Some limitations of this review should be mentioned. The results of this narrative review suggest that anatomically non-representative femora may affect the validity of biomechanical testing of osteosynthesis. Although there are no known clinical failures directly attributable to the use of non-representative bone surrogates in biomechanical testing, the important conclusion of this paper is that discrepancies do exist and that those performing biomechanical testing should be aware of the differences from human dimensions and characteristics.

Furthermore, in this study only the measurements given by the manufacturer were considered. Other parameters, such as femoral curvature or bowing, which may have an impact on the placement of intramedullary nails, were therefore not considered.

Moreover, only the relevance of the absence of population variability in femoral bone surrogates for osteosynthesis was discussed. However, the non-representative dimensions could also have an impact on other biomechanical tests, e.g. in knee or hip arthroplasty.

Finally, only the femur was considered in this review. However, the concerns raised regarding the lack of population variability in bone surrogates for biomechanical testing of osteosynthesis are certainly transferable to other anatomical regions.

Conclusions

In conclusion, the Sawbones femur models represent a limited part of the general population. Therefore, when planning a biomechanical study, bone density should be considered in addition to the choice of model size and whether suboptimal patient care could occur due to the lack of integration of population variability in the artificial bone.