Heidari et al 2020a - Revision history

Rimni: /* 2.3 Boundary conditions and loading */

2021-02-12T13:16:49Z

‎2.3 Boundary conditions and loading

Rimni: /* 2.2 Modeling of Implant components */

2021-02-12T12:08:05Z

‎2.2 Modeling of Implant components

Rimni at 12:04, 12 February 2021

2021-02-12T12:04:42Z

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Scipediacontent at 10:45, 27 March 2020

2020-03-27T10:45:33Z

Scipediacontent at 10:44, 27 March 2020

2020-03-27T10:44:09Z

Scipediacontent at 10:42, 27 March 2020

2020-03-27T10:42:34Z

Scipediacontent at 10:35, 27 March 2020

2020-03-27T10:35:22Z

Scipediacontent at 09:48, 27 March 2020

2020-03-27T09:48:32Z

Rimni at 09:35, 27 March 2020

2020-03-27T09:35:38Z

Rimni at 09:34, 27 March 2020

2020-03-27T09:34:04Z

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 ! Material !! Elastic module (GPa) !! Poison ratio
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 |  style="text-align: center;"|Humeral/Ulnar component (Titanium)
 |  style="text-align: center;"|110
 |  style="text-align: center;"|0.3
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 |  style="text-align: center;"|Bone
 |  style="text-align: center;"|16.7
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 |  style="text-align: center;"|Cement

← Older revision		Revision as of 12:08, 12 February 2021
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−	The anterior flange transfers a portion of the load through the elbow prosthesis stem to the distal region of the humerus, which diminishes the probability of humeral loosening [27,28]. Therefore, the flange was modeled with various shapes including concave, trapezium and rectangular. CT data of distal humerus was used to determine the curve angle of the trapezium model. The angle of where the flange faces distal humerus was nearly 7.3°. These three models were analyzed under the same cross-section (rectangular), the same humeral stem length (150mm), and similar dimensions (Figure 8). Bone graft was used between the distal humerus and flange. AP compression force (1N) and axial torsion (1N.cm) were applied to the articular surface of the humeral component.	+	The anterior flange transfers a portion of the load through the elbow prosthesis stem to the distal region of the humerus, which diminishes the probability of humeral loosening [27,28]. Therefore, the flange was modeled with various shapes including concave, trapezium and rectangular. CT data of distal humerus was used to determine the curve angle of the trapezium model. The angle of where the flange faces distal humerus was nearly 7.3°. These three models were analyzed under the same cross-section (rectangular), the same humeral stem length (150mm), and similar dimensions ([[#img-8\|Figure 8]]). Bone graft was used between the distal humerus and flange. AP compression force (1N) and axial torsion (1N.cm) were applied to the articular surface of the humeral component.

	{\| style="text-align: center; border: 1px solid #BBB; margin: 1em auto; width: 80%;"		{\| style="text-align: center; border: 1px solid #BBB; margin: 1em auto; width: 80%;"

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 Figure 20 presents the stress distribution in the implanted and intact ulna (intact ulna acts as a reference) under AP compression. It was observed the stress distribution in diaphyseal and distal regions. The proximal part experienced stress shielding. In anterior of the proximal area, intact ulna experienced stress at 70.68 kPa. The Von Mises stress distribution in proximal of the implanted ulna is shown in Figure 21 (a). The lowest stress was seen in width concave and all sides concave <math>R=10</math>mm stems (61.9 kPa), while the highest stress was observed in the diamond stem at 63.9 kPa. Concave stem (<math>R=8</math>) showed the second highest stress at 62.9 kPa. All sides concave <math>R=12</math>mm and <math>R=10</math>mm stems experienced stress at the same amount at 62.8 kPa, while the rectangular stem presented at 62.2 kPa. The effect of cross-section on the stress shielding (under AP compression force) in the ulnar stem is Figure 21 (b). Moderate growth in stress shielding was seen from 6.2% for all sides concave <math>R=8</math>mm stem to nearly 6.4% for convex <math>R=10</math>mm stem which presented the highest amount than all models. The lowest stress shielding was registered for the diamond stem at 6.1%. The similar stress shielding was observed for width convex and rectangular stems at approximately 6.3%.
-{| style="text-align: center; border: 1px solid #BBB; margin: 1em auto; width:60%;"
+{| style="text-align: center; border: 1px solid #BBB; margin: 1em auto; width: auto;max-width: auto;"
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 Figure 19 (a) presents average stress in the distal region of the implanted humerus for different cross-sections. The diamond stem showed the highest stress at 66.2 kPa, while all sides convex and rectangular stems experienced the least micro-motion at 60.5 kPa. Triangular stem experienced less stress than the diamond stem (64.45 kPa). Two sides triangular (1mm) and (0.5mm) stems presented higher stress compared to convex and concave at 61.8 kPa and 63 kPa respectively. The stress shielding for the humeral stem with different cross-sections is shown in Figure 19 (b). Triangular (6.0%) and diamond (5.8%) stems presented the lowest stress shielding. Approximately, it was seen the same stress shielding at 6.4% for concaves <math>R=50</math>mm, <math>R=70</math>mm, rectangular and sides concave stems. Concave stems experienced slightly less stress shielding than convex stems. However, two sides triangular (1mm), at 6.2%.stem, presented moderately less than two sides triangular (0.5mm) stem at 6.4%.
-{| style="text-align: center; border: 1px solid #BBB; margin: 1em auto; width: 40;"
+{| style="text-align: center; border: 1px solid #BBB; margin: 1em auto; width: auto;max-width: auto;"
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-{| style="text-align: center; border: 1px solid #BBB; margin: 1em auto; width: 40;"
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 Figure 18 presents the Von Mises stress distribution in the implanted and intact humerus (intact humerus acts as a reference) under AP compression. The humerus was assigned into three regions of interest (ROI): distal, diaphyseal, and proximal. Reduction of stress expresses a decrease reduction in the load and consequently the growth of the stress shielding. No effective changes were seen for stress distribution between implanted and intact humerus in diaphyseal and proximal areas. Stress shielding occurred in the distal region. The anterior of the distal area showed stress at 72.90 kPa in the intact humerus.
-{| style="text-align: center; border: 1px solid #BBB; margin: 1em auto; width: 40;"
 |-
-|style="padding:10px;"| [[Image:Heidari_et_al_2020a_7520_Figura18.png|center|600px]]
+|style="padding:10px;"| [[Image:Review_392361619556-image58.jpeg|268px]]
 |- style="text-align: center; font-size: 75%;"
 | colspan="2" style="padding:10px;"| '''Figure 18'''. Von Mises stress in intact humerus bone (left). Implanted humerus bone with rectangular stem cross-section (right)

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 ===2.1 Modeling of elbow bones===
-Mimics software (Ver.10, Materialise NV., Belgium), through computed tomography (CT), was used to construct elbow bones. CT data set involved the slice thickness of 0.7mm and the resolution of 0.477mm (Figure 1). Then, the constructed elbow bones were imported into SolidWorks. In this study, SolidWorks software was used for simulated surgery and modeling all components of the linked elbow implant. Simulated surgery of humerus and ulna bones was carried out to insert modeled implant into the constructed bone. Therefore, a hole was made in humerus/ulna bones and the implant left these bones with a hole. The initial step to prepare of the humerus was bone resection from the midportion of the capitellum and trochlea (Figure 2). As a result, the center axis of flexion-extension was identified. If the capitellum is assumed a circle and sighted from the lateral position of the elbow, the center axis of rotation is placed in the center of the circle. A similar procedure was carried out to prepare ulna (Figure 3). The ulnar component was inserted into the ulna to ensure a good fit before cementing. Conventional cemented fixation was the basis of modeling of all stems therefore, they had smooth surfaces. The optimum thickness of cement (between 1.0 and 1.5 mm) was considered to avoid high failure in the cement-bone [26].
+Mimics software (Ver.10, Materialise NV., Belgium), through computed tomography (CT), was used to construct elbow bones. CT data set involved the slice thickness of 0.7mm and the resolution of 0.477mm (Figure 1). Then, the constructed elbow bones were imported into SolidWorks. In this study, SolidWorks software was used for simulated surgery and modeling all components of the linked elbow implant. Simulated surgery of humerus and ulna bones was carried out to insert modeled implant into the constructed bone. Therefore, a hole was made in humerus/ulna bones and the implant left these bones with a hole. The initial step to prepare of the humerus was bone resection from the midportion of the capitellum and trochlea (Figure 2). As a result, the center axis of flexion-extension was identified. If the capitellum is assumed a circle and sighted from the lateral position of the elbow, the center axis of rotation is placed in the center of the circle. A similar procedure was carried out to prepare ulna (Figure 3). The ulnar component was inserted into the ulna to ensure a good fit before cementing. Conventional cemented fixation was the basis of modeling of all stems therefore, they had smooth surfaces. The optimum thickness of cement (between 1.0 and 1.5mm) was considered to avoid high failure in the cement-bone [26].
 {| style="text-align: center; border: 1px solid #BBB; margin: 1em auto; width: 40%;"
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 ===2.2 Modeling of Implant components===
-Many studies have carried out to analyze the ideal stem cross-sectional to control implant loosening. Dimension for modeling elbow implant was obtained from a real Coonrad-Morrey implant. Various shapes were modeled to modify a proper cross-section for the humeral component. The cross-sections included diamond, rectangular, triangular, width triangular, concave, and convex. The length of the humeral stem was the same for all cross-sections at 150 mm (Figure 4). Several sub-designs were noticed for the concave cross-section including width concave with variant radiuses (25mm to 70mm) and all sides concave. In the same way, different sub-designs were assumed for convex cross-section, including width convex with the different radiuses (25mm to 70mm) and all sides convex (Figure 5). Furthermore, various cross-sections including concave, rectangular, convex, diamond modeled to develop a sufficient ulnar component (Figure 6). Several sub-designs were used for concave cross-section including height concave, all sides concave, and width concave (radius from 8mm to 12mm). Similarly, different sub-designs were considered for convex cross-section including width convex with various (radius of 8mm to 12mm), all sides convex, and height concave (Figure 7).
+Many studies have carried out to analyze the ideal stem cross-sectional to control implant loosening. Dimension for modeling elbow implant was obtained from a real Coonrad-Morrey implant. Various shapes were modeled to modify a proper cross-section for the humeral component. The cross-sections included diamond, rectangular, triangular, width triangular, concave, and convex. The length of the humeral stem was the same for all cross-sections at 150mm (Figure 4). Several sub-designs were noticed for the concave cross-section including width concave with variant radiuses (25mm to 70mm) and all sides concave. In the same way, different sub-designs were assumed for convex cross-section, including width convex with the different radiuses (25mm to 70mm) and all sides convex (Figure 5). Furthermore, various cross-sections including concave, rectangular, convex, diamond modeled to develop a sufficient ulnar component (Figure 6). Several sub-designs were used for concave cross-section including height concave, all sides concave, and width concave (radius from 8mm to 12mm). Similarly, different sub-designs were considered for convex cross-section including width convex with various (radius of 8mm to 12mm), all sides convex, and height concave (Figure 7).
 {| style="text-align: center; border: 1px solid #BBB; margin: 1em auto; width: 50%;"
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-The anterior flange transfers a portion of the load through the elbow prosthesis stem to the distal region of the humerus, which diminishes the probability of humeral loosening [27,28]. Therefore, the flange was modeled with various shapes including concave, trapezium and rectangular. CT data of distal humerus was used to determine the curve angle of the trapezium model. The angle of where the flange faces distal humerus was nearly 7.3°. These three models were analyzed under the same cross-section (rectangular), the same humeral stem length (150 mm), and similar dimensions (Figure 8). Bone graft was used between the distal humerus and flange. AP compression force (1N) and axial torsion (1N.cm) were applied to the articular surface of the humeral component.
+The anterior flange transfers a portion of the load through the elbow prosthesis stem to the distal region of the humerus, which diminishes the probability of humeral loosening [27,28]. Therefore, the flange was modeled with various shapes including concave, trapezium and rectangular. CT data of distal humerus was used to determine the curve angle of the trapezium model. The angle of where the flange faces distal humerus was nearly 7.3°. These three models were analyzed under the same cross-section (rectangular), the same humeral stem length (150mm), and similar dimensions (Figure 8). Bone graft was used between the distal humerus and flange. AP compression force (1N) and axial torsion (1N.cm) were applied to the articular surface of the humeral component.
 {| style="text-align: center; border: 1px solid #BBB; margin: 1em auto; width: 80%;"
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 ===2.4 Mesh convergence===
-An ulnar stem was used to investigate mesh convergence. For this purpose, a series of runs were accomplished with similar conditions except for the size of the element, which was reduced between each run. Tetrahedron element sizes from 4 mm to 0.5 mm were utilized to mesh model (Figure 11). Average Von Mises stress in the ulnar stem is shown in Figure 12.  It was seen a considerable diversity between the refined models and coarsest ones. The mesh sizes of 1 mm (41.14 kPa), 0.75 mm (41.15 kPa) and 0.5 mm (41.19 kPa) presented approximately the same stresses. Undoubtedly, more time was spent running the analysis with the mesh size of 0.5 mm and 0.75 mm than the size of 1 mm while, the results were similar. Hence, convergence was reached when there was no qualitative change in the stresses. Consequently, this study was used a mesh size of 1 mm.
+An ulnar stem was used to investigate mesh convergence. For this purpose, a series of runs were accomplished with similar conditions except for the size of the element, which was reduced between each run. Tetrahedron element sizes from 4mm to 0.5mm were utilized to mesh model (Figure 11). Average Von Mises stress in the ulnar stem is shown in Figure 12.  It was seen a considerable diversity between the refined models and coarsest ones. The mesh sizes of 1mm (41.14 kPa), 0.75mm (41.15 kPa) and 0.5mm (41.19 kPa) presented approximately the same stresses. Undoubtedly, more time was spent running the analysis with the mesh size of 0.5mm and 0.75mm than the size of 1mm while, the results were similar. Hence, convergence was reached when there was no qualitative change in the stresses. Consequently, this study was used a mesh size of 1mm.
 {| style="text-align: center; border: 1px solid #BBB; margin: 1em auto; width: auto;max-width: auto;"
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 |style="padding:10px;"| [[File:Review_392361619556_6297_19.jpg]]
 |- style="text-align: center; font-size: 75%;"
-| colspan="1" style="padding-bottom:10px;"| '''Figure 13'''. Micro-motion for different humeral stem cross-sections under AP compression force. (a) Bone-cement interface. (b) Cement-implant interface: (a) Rectangular. (b) Diamond.(c) Triangular. (d) Concave <math> R=25</math>mm. (e) Concave <math> R=50</math>mm. (f) Concave <math>R=70</math>mm. (g) All sides concave. (h) Two sides triangular (1mm). (i) Two sides triangular (0.5 mm).  (j) Convex <math>R=25</math>mm. (k) Convex <math> R=50</math>mm. (l) Convex <math>R=70</math>mm. (m) All sides convex
+| colspan="1" style="padding-bottom:10px;"| '''Figure 13'''. Micro-motion for different humeral stem cross-sections under AP compression force. (a) Bone-cement interface. (b) Cement-implant interface: (a) Rectangular. (b) Diamond.(c) Triangular. (d) Concave <math> R=25</math>mm. (e) Concave <math> R=50</math>mm. (f) Concave <math>R=70</math>mm. (g) All sides concave. (h) Two sides triangular (1mm). (i) Two sides triangular (0.5mm).  (j) Convex <math>R=25</math>mm. (k) Convex <math> R=50</math>mm. (l) Convex <math>R=70</math>mm. (m) All sides convex
 |}
-The effect of cross-section on micro-motion of the humeral stem (under torsion force) is shown in Figure 14. In the cement-bone interface, diamond (<math>34\,\mu</math>m) and triangular (<math>90\,\mu</math>m) stems showed the highest and the lowest micro-motions respectively. Two sides triangular (0.5 mm) and concave <math> R=50</math>mm stems presented the same micro-motion at <math>74\,\mu</math>m, which was considerably higher than stems with concave <math> R=50</math>mm, triangular, convex <math> R=25</math>, convex <math> R=50</math>, and convex <math>R=70</math> mm cross-sections. In the implant-cement interface, the smallest micro-motion was observed in diamond and convex <math> R=50</math>mm stems at <math>52\,\mu</math>m and <math>42\,\mu</math>m respectively while rectangular and all sides concave experienced micro-motion at <math>103\,\mu</math>m. Figures were slightly increased from convex <math>R=70</math> (<math>54\,\mu</math>m) to <math>89\,\mu</math>m for the two sides triangular (1mm). All sides concave stem showed higher micro-motion (<math>89\,\mu</math>m) than concave <math>R=70</math>mm and <math> R=50</math>mm at <math>81\,\mu</math>m.
+The effect of cross-section on micro-motion of the humeral stem (under torsion force) is shown in Figure 14. In the cement-bone interface, diamond (<math>34\,\mu</math>m) and triangular (<math>90\,\mu</math>m) stems showed the highest and the lowest micro-motions respectively. Two sides triangular (0.5mm) and concave <math> R=50</math>mm stems presented the same micro-motion at <math>74\,\mu</math>m, which was considerably higher than stems with concave <math> R=50</math>mm, triangular, convex <math> R=25</math>, convex <math> R=50</math>, and convex <math>R=70</math>mm cross-sections. In the implant-cement interface, the smallest micro-motion was observed in diamond and convex <math> R=50</math>mm stems at <math>52\,\mu</math>m and <math>42\,\mu</math>m respectively while rectangular and all sides concave experienced micro-motion at <math>103\,\mu</math>m. Figures were slightly increased from convex <math>R=70</math> (<math>54\,\mu</math>m) to <math>89\,\mu</math>m for the two sides triangular (1mm). All sides concave stem showed higher micro-motion (<math>89\,\mu</math>m) than concave <math>R=70</math>mm and <math> R=50</math>mm at <math>81\,\mu</math>m.
 {| style="text-align: center; border: 1px solid #BBB; margin: 1em auto; width: auto;max-width: auto;"
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 |style="padding:10px;"| [[File:Review_392361619556_7283_20.jpg]]
 |- style="text-align: center; font-size: 75%;"
-| colspan="1" style="padding-bottom:10px;"| '''Figure 14'''. Micro-motion for different humeral stem cross-sections under torsion force in (a) Bone-cement interface. (b) Cement-implant interface: (a) Rectangular. (b) Diamond. (c) Triangular. (d) Concave <math> R=25</math>mm. (e) Concave <math> R=50</math>mm. (f) Concave <math>R=70</math>mm. (g) All sides concave. (h) Two sides triangular (1mm). (i) Two sides triangular (0.5 mm).  (j) Convex <math>R=25</math>mm. (k) Convex <math> R=50</math>mm. (l) Convex <math>R=70</math>mm. (m) All sides convex
+| colspan="1" style="padding-bottom:10px;"| '''Figure 14'''. Micro-motion for different humeral stem cross-sections under torsion force in (a) Bone-cement interface. (b) Cement-implant interface: (a) Rectangular. (b) Diamond. (c) Triangular. (d) Concave <math> R=25</math>mm. (e) Concave <math> R=50</math>mm. (f) Concave <math>R=70</math>mm. (g) All sides concave. (h) Two sides triangular (1mm). (i) Two sides triangular (0.5mm).  (j) Convex <math>R=25</math>mm. (k) Convex <math> R=50</math>mm. (l) Convex <math>R=70</math>mm. (m) All sides convex
 |}
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 The flange protects the implant-bone-cement interfaces and affects mechanical factors and consequently reduction of implant loosening. Herren et al. studied the influences of an anterior flange to the GSB III humeral component [52]. No significant diversities were observed in displacements or the stress distribution, indicating the flange had a small influence on this design. Quenneville et al. studied the impact of a flange on strains in the distal region of the humerus. It was seen that strains are affected by a flange using different geometric [53]. In a similar way, our study proved that micro-motion is affected by the change of flange shape.
-Our results showed that a concave flange transfers the lowest micro-motion to the implant-bone-cement interface compared with rectangular and trapezium. Theoretically, the flange increases the rotational stability of the humeral component. The flange affects the transmission of micro-motion to the distal region of humeral stem. Increased micro-motion leads to more movement between surfaces, which widen the boundaries in interfaces and the migration of implant components. Clinically, this migration is seen as radiolucency surrounding the stem. If radiolucency is increased to more than 2 mm, loosening probably occurs. Hence, less micro-motion transferred by the concave flange to interfaces may cause lower loosening. Although the effect of the introduced flange in this study, concave, on micro-motion is little, it probably decreases loosening in a long time. The concave flange may decrease more stress shielding in distal humerus bone rather than other flanges.
+Our results showed that a concave flange transfers the lowest micro-motion to the implant-bone-cement interface compared with rectangular and trapezium. Theoretically, the flange increases the rotational stability of the humeral component. The flange affects the transmission of micro-motion to the distal region of humeral stem. Increased micro-motion leads to more movement between surfaces, which widen the boundaries in interfaces and the migration of implant components. Clinically, this migration is seen as radiolucency surrounding the stem. If radiolucency is increased to more than 2mm, loosening probably occurs. Hence, less micro-motion transferred by the concave flange to interfaces may cause lower loosening. Although the effect of the introduced flange in this study, concave, on micro-motion is little, it probably decreases loosening in a long time. The concave flange may decrease more stress shielding in distal humerus bone rather than other flanges.
 Several limitations were observed in this study. Only one elbow sample was considered for this study. Considering a wide range of elbow bones with different sizes may deliver more accurate results. Furthermore, considering complex loading in the implanted elbow may present more reliable results. It is hoped that future studies may be able to accommodate these deficiencies which may or may not play an important role in the more accurate design of elbow implants. Many different recommendations can be stated for future extensions of the introduced approach in this study. Investigating the effect of changing the flexion-extension axis in micro-motion transferred to the bone-cement-implant interfaces can develop the approach further. Furthermore, using the same approach, remodeling of the humerus and ulna bones can be investigated after surgery. As well as, future work can concern the kinematic of intact and elbow replaced by the modified implant.