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Dynamic muscle loading and mechanotransduction
Bone 51 (2012) 826–827
Contents lists available at SciVerse ScienceDirect
Bone journal homepage: www.elsevier.com/locate/bone
Commentary
Dynamic muscle loading and mechanotransduction
The beneficial effects of exercise on skeletal health have been known for many years. Experimental data regarding the physical and biological processes that control bone adaptation reveal new avenues of therapy, which have the potential to augment bone mass and strength significantly. One such area is the manipulation of bone's physical environment by unconventional methods (e.g., high-frequency vibration and ultrasound). These unconventional methods permit some of the bonebuilding effects of mechanotransduction to occur, while preventing some of the detrimental effects of conventional mechanical loading (e.g., fracture or soft tissue injury from high strain application). In this issue of Bone, Hu et al. [1] present a novel, unconventional technique for applying mechanical loading to bone, with the goal of enhancing bone formation and/or preventing bone loss. This loading scheme, referred to as dynamic hydraulic stimulation (DHS), applies noninvasive dynamic pressure (30 mm Hg static + 30 mm Hg dynamic) to the skin and muscle tissues that surround the tibia of a hindlimbsuspended rat for 20 min per day, 5 days per week. After 4 weeks of DHS treatment, bone volume fraction and bone formation rate were increased by 83% and 190%, respectively, compared to their nonstimulated, tail-suspended counterparts. Two unique features of this non-invasive loading method are the stimulation of bone formation through muscle compression and the application of lateral loads at low frequency. Qin et al. [2] previously reported increased bone formation in response to direct intramedullary pressure (ImP) applied to an isolated turkey ulna. The same group also reported that electric stimulation of the quadriceps muscle elevates ImP and increases bone formation in the femur [3]. In this report, DHS additionally applies mechanical stimulation through the gastrocnemius and other muscles around the tibia. Instead of applying axial loads, which are more closely linked to routine physical activities, DHS employs lateral loads. Application of lateral loads is also used in a joint-loading modality (JL) such as knee loading and ankle loading [4]. It is reported that JL to a mouse knee alters ImP [5]. DHS applies loads to the diaphysis by constricting circumferentially, while JL applies loads to the epiphysis and metaphysis by compressing at two confined regions. Although the mechanism of action underlying enhanced bone formation by DHS might not represent natural adaptation of bone through physical activity, this modality not only offers a potential strategy for strengthening bone in patients with osteoporosis, but also provides a model system for dissecting mechanotransduction of bone. The working hypothesis of DHS' mechanism of action consists of three major steps: alterations in ImP, induction of interstitial fluid flow (IFF), and mechanosensing by osteocytes followed by activation of Wnt signaling in osteoblasts. In the first step, constrictive loads might induce a volumetric change in the intramedullary space, and this change may result in an ImP increase. A small amount of strain, in the order of 10 μstrain in the diaphysis, might be the cause of 8756-3282/$ – see front matter © 2012 Elsevier Inc. All rights reserved. http://dx.doi.org/10.1016/j.bone.2012.07.025
this pressure change. The effect of strain is likely to be local at the site of loading; however, the effect of the increased ImP is not confined to the loading site but is distributed throughout the intramedullary cavity. This first step seems consistent with a series of experiments conducted by the Frangos group. These experiments correlate bone formation to enhanced ImP using an implanted miniature pump that cyclically modulated pressure [6] and partial femoral vein ligation in hindlimb-suspended rats [7]. However, in the case of vein ligation, it is not clear that the change in ImP is the predominant factor in counteracting bone loss as many other conditions are affected by this procedure. It is also unclear whether temporal alteration of ImP is sufficient to induce mechanosensing, since cells are in general more sensitive to shear stress than normal stress. Therefore, in the second step, a mechanism for converting a temporal change in ImP into a spatial gradient of ImP is proposed. On the one hand, in order to elevate ImP, a bone cavity must be liquid-tight; on the other hand, to induce a special gradient of ImP, fluid in the cavity needs to flow somewhere and restrictions should be relaxed. The bone cavity consists of a mixture of fluid and solid materials, and the question is how liquid-tight and permeable the bone cavity should be in order to generate a temporal change in ImP followed by a special gradient of ImP. One of the important parameters to answer this question is conductivity as a function of Darcy's permeability (m2). This conductivity dictates pressure-driven flow induction, and it depends on the bone matrix porosities, including collagen–apatite porosity, the lacunar–canalicular porosity, and the vascular porosity [8]. A layer of tissues that surrounds bone is also important, since it provides a semipermeable boundary layer characterized by another kind of permeability (m/s), defined by the thickness of the layer and a diffusion coefficient. Porous structures in bone and the surrounding layer may all be connected and their network as a whole may define temporal and spatial distributions of ImP. The third step is the transition from IFF to load-driven molecular events represented by Wnt signaling [9]. An energy level of thermal noise is estimated in the order of kT (4 × 10 −21 J), where k is the Boltzmann constant and T is the body absolute temperature. In order to induce a meaningful mechanical stimulation, one assumption is that IFF should be able to generate the amount of energy equivalent to one ATP molecule or ~ 10kT (the energy generated by one ATP molecule is approximately 10 times larger than kT). A rough estimation of this concept is as follows: in order to move part of a mechanosensing molecule by several nanometers, 10kT of energy can generate a force of ~ 10 pN. This amount of force corresponds to a surface area of 10 μm 2 under shear stress at 1 Pa (10 dyn/cm 2). In order for a single osteocyte process (~ 20 μm long, 50–400 nm in diameter) [10] to receive 10 pN and move a molecular sensor by several nanometers, it is plausible that there is an amplification mechanism
Commentary
such as the one proposed by Weinbaum [11]. Alternatively, local fluctuations from anisotropic and stochastic dynamics may induce a spectrum of stress distributions that might facilitate activation of molecular interactions [12]. Besides the above-mentioned working hypothesis, DHS allows us to consider other mechanisms indirectly linked to mechanical stimulation. Mechanical loading must trigger migration of osteoblasts to the endosteum and periosteum. Because ImP is altered in the intramedullary cavity, it is interesting to ask whether alterations in ImP accelerate migration of osteoblasts and whether any differences exist in the rate of bone formation in the endosteum and periosteum. An additional question is whether bone formation would take place equally throughout the length of the tibia in cortical and trabecular bone or whether any preferential location exists. Lastly, an intriguing question is the importance of local vs. systemic effects in connection to blood circulation and neuronal signaling [7,13]. DHS seems to provide a promising model system for evaluating the role of blood circulation and neuronal signaling in load-driven bone formation. In summary, as the field continues to make advances in our understanding of the physical and structural conditions that stimulate bone formation, safer and perhaps more effective approaches to loadinduced bone formation can be achieved. These developments could not be more timely, as physicians look for alternative means to maintain bone mass in patients once they complete an osteoporosis drug treatment regimen. Several recently-discovered therapies (e.g., Rank-L inhibitors, PTH fragments, and cathepsin K inhibitors) appear to have little to no residual effects once treatment is stopped because they are not stored in the bone tissues in active form. Therefore, bone density is likely to return to baseline once these next-generation therapies are stopped. As the more long-lasting bisphosphonate therapies fall out of favor and newly generated compounds rise in popularity among physicians, the need to maintain bone mass after treatment withdrawal will become more urgent. Therefore, a safe, effective mechanical loading modality potentially provides an alternative treatment for maintaining BMD without inducing deleterious side effects. Although DHS has not been tested clinically, it is not difficult to envision its implementation via a portable loading device that may resemble the inflatable cuff of a blood pressure monitoring system. Its extension from diaphysis to the femoral neck and head seems natural, but its application to spine treatment is less evident. It would be interesting to explore DHS' potential capability not only to enhance bone formation beyond the immediate local environment, but also to stimulate fracture healing without direct application of mechanical loads to the fracture site.
827
[4] Zhang P, Turner CH, Yokota H. Joint loading-driven bone formation and signaling pathways predicted from genome-wide expression profiles. Bone 2009;44:989-98. [5] Zhang P, Su M, Liu Y, Hsu A, Yokota H. Knee loading dynamically alters intramedullary pressure in mouse femora. Bone 2007;40:538-43. [6] Kwon RY, Meays DR, Tang WJ, Frangos JA. Microfluidic enhancement of intramedullary pressure increases interstitial fluid flow and inhibits bone loss in hindlimb suspended mice. J Bone Miner Res 2010;25:1798-807. [7] Bergula AP, Huang W, Frangos JA. Femoral vein ligation increases bone mass in the hindlimb suspended rat. Bone 1999;24:171-7. [8] Fritton SP, Weinbaum S. Fluid and solute transport in bone: flow-induced mechanotransduction. Annu Rev Fluid Mech 2009;41:347-74. [9] Sawakami K, Robling AG, Ai M, Pitner ND, Liu D, Warden SJ, et al. The Wnt co-receptor LRP5 is essential for skeletal mechanotransduction but not for the anabolic bone response to parathyroid hormone treatment. J Biol Chem 2006;281: 23698-711. [10] Sugawara Y, Kamioka H, Honjo T, Tezuka K, Takano-Yamamoto T. Three-dimensional reconstruction of chick calvarial osteocytes and their cell processes using confocal microscopy. Bone 2005;36:877-83. [11] Han Y, Cowen SC, Schaffler MB, Weinbaum S. Mechanotransduction and strain amplification in osteocyte cell processes. Proc Natl Acad Sci U S A 2004;101: 16689-94. [12] Knothe Tate ML. Top down and bottom up engineering of bone. J Biomech 2011;44:304-12. [13] Sample SJ, Behan M, Smith L, Oldenhoff WE, Markel MD, Kalscheur VL, et al. Functional adaptation to loading of a single bone is neuronally regulated and involves multiple bones. J Bone Miner Res 2008;23:1372-81.
Hiroki Yokota Biomechanics and Biomaterials Research Center, Indiana University Purdue University Indianapolis, Indianapolis, IN 46202, USA Department of Biomedical Engineering, Indiana University Purdue University Indianapolis, Indianapolis, IN 46202, USA Department of Anatomy and Cell Biology, Indiana University School of Medicine, Indianapolis, IN 46202, USA Corresponding author at: Biomechanics and Biomaterials Research Center, Department of Biomedical Engineering, Indiana University Purdue University Indianapolis, 723 West Michigan Street, Indianapolis, IN 46202, USA. Fax: +1 317 278 5244. E-mail address: [email protected]. Andrés Tovar Biomechanics and Biomaterials Research Center, Indiana University Purdue University Indianapolis, Indianapolis, IN 46202, USA Department of Mechanical Engineering, Indiana University Purdue University Indianapolis, Indianapolis, IN 46202, USA
Acknowledgment The work is in part supported by NIH R01AR052144 (HY) and R01AR053237 (AR). References [1] Hu M, Cheng J, Qin YX. Dynamic hydraulic flow stimulation on mitigation of trabecular bone loss in a rat functional disuse model. Bone 2012;51:819-25. [2] Qin YX, Kaplan T, Saldanha A, Rubin C. Fluid pressure gradients, arising from oscillations in intramedullary pressure, is correlated with the formation of bone and inhibition of intracortical porosity. J Biomech 2003;36:1427-37. [3] Qin YX, Lam H. Intramedullary pressure and matrix strain induced by oscillatory skeletal muscle stimulation and its potential in adaptation. J Biomech 2009;42: 140-5.
Alexander Robling Biomechanics and Biomaterials Research Center, Indiana University Purdue University Indianapolis, Indianapolis, IN 46202, USA Department of Biomedical Engineering, Indiana University Purdue University Indianapolis, Indianapolis, IN 46202, USA Department of Anatomy and Cell Biology, Indiana University School of Medicine, Indianapolis, IN 46202, USA 8 June 2012
Contents lists available at SciVerse ScienceDirect
Bone journal homepage: www.elsevier.com/locate/bone
Commentary
Dynamic muscle loading and mechanotransduction
The beneficial effects of exercise on skeletal health have been known for many years. Experimental data regarding the physical and biological processes that control bone adaptation reveal new avenues of therapy, which have the potential to augment bone mass and strength significantly. One such area is the manipulation of bone's physical environment by unconventional methods (e.g., high-frequency vibration and ultrasound). These unconventional methods permit some of the bonebuilding effects of mechanotransduction to occur, while preventing some of the detrimental effects of conventional mechanical loading (e.g., fracture or soft tissue injury from high strain application). In this issue of Bone, Hu et al. [1] present a novel, unconventional technique for applying mechanical loading to bone, with the goal of enhancing bone formation and/or preventing bone loss. This loading scheme, referred to as dynamic hydraulic stimulation (DHS), applies noninvasive dynamic pressure (30 mm Hg static + 30 mm Hg dynamic) to the skin and muscle tissues that surround the tibia of a hindlimbsuspended rat for 20 min per day, 5 days per week. After 4 weeks of DHS treatment, bone volume fraction and bone formation rate were increased by 83% and 190%, respectively, compared to their nonstimulated, tail-suspended counterparts. Two unique features of this non-invasive loading method are the stimulation of bone formation through muscle compression and the application of lateral loads at low frequency. Qin et al. [2] previously reported increased bone formation in response to direct intramedullary pressure (ImP) applied to an isolated turkey ulna. The same group also reported that electric stimulation of the quadriceps muscle elevates ImP and increases bone formation in the femur [3]. In this report, DHS additionally applies mechanical stimulation through the gastrocnemius and other muscles around the tibia. Instead of applying axial loads, which are more closely linked to routine physical activities, DHS employs lateral loads. Application of lateral loads is also used in a joint-loading modality (JL) such as knee loading and ankle loading [4]. It is reported that JL to a mouse knee alters ImP [5]. DHS applies loads to the diaphysis by constricting circumferentially, while JL applies loads to the epiphysis and metaphysis by compressing at two confined regions. Although the mechanism of action underlying enhanced bone formation by DHS might not represent natural adaptation of bone through physical activity, this modality not only offers a potential strategy for strengthening bone in patients with osteoporosis, but also provides a model system for dissecting mechanotransduction of bone. The working hypothesis of DHS' mechanism of action consists of three major steps: alterations in ImP, induction of interstitial fluid flow (IFF), and mechanosensing by osteocytes followed by activation of Wnt signaling in osteoblasts. In the first step, constrictive loads might induce a volumetric change in the intramedullary space, and this change may result in an ImP increase. A small amount of strain, in the order of 10 μstrain in the diaphysis, might be the cause of 8756-3282/$ – see front matter © 2012 Elsevier Inc. All rights reserved. http://dx.doi.org/10.1016/j.bone.2012.07.025
this pressure change. The effect of strain is likely to be local at the site of loading; however, the effect of the increased ImP is not confined to the loading site but is distributed throughout the intramedullary cavity. This first step seems consistent with a series of experiments conducted by the Frangos group. These experiments correlate bone formation to enhanced ImP using an implanted miniature pump that cyclically modulated pressure [6] and partial femoral vein ligation in hindlimb-suspended rats [7]. However, in the case of vein ligation, it is not clear that the change in ImP is the predominant factor in counteracting bone loss as many other conditions are affected by this procedure. It is also unclear whether temporal alteration of ImP is sufficient to induce mechanosensing, since cells are in general more sensitive to shear stress than normal stress. Therefore, in the second step, a mechanism for converting a temporal change in ImP into a spatial gradient of ImP is proposed. On the one hand, in order to elevate ImP, a bone cavity must be liquid-tight; on the other hand, to induce a special gradient of ImP, fluid in the cavity needs to flow somewhere and restrictions should be relaxed. The bone cavity consists of a mixture of fluid and solid materials, and the question is how liquid-tight and permeable the bone cavity should be in order to generate a temporal change in ImP followed by a special gradient of ImP. One of the important parameters to answer this question is conductivity as a function of Darcy's permeability (m2). This conductivity dictates pressure-driven flow induction, and it depends on the bone matrix porosities, including collagen–apatite porosity, the lacunar–canalicular porosity, and the vascular porosity [8]. A layer of tissues that surrounds bone is also important, since it provides a semipermeable boundary layer characterized by another kind of permeability (m/s), defined by the thickness of the layer and a diffusion coefficient. Porous structures in bone and the surrounding layer may all be connected and their network as a whole may define temporal and spatial distributions of ImP. The third step is the transition from IFF to load-driven molecular events represented by Wnt signaling [9]. An energy level of thermal noise is estimated in the order of kT (4 × 10 −21 J), where k is the Boltzmann constant and T is the body absolute temperature. In order to induce a meaningful mechanical stimulation, one assumption is that IFF should be able to generate the amount of energy equivalent to one ATP molecule or ~ 10kT (the energy generated by one ATP molecule is approximately 10 times larger than kT). A rough estimation of this concept is as follows: in order to move part of a mechanosensing molecule by several nanometers, 10kT of energy can generate a force of ~ 10 pN. This amount of force corresponds to a surface area of 10 μm 2 under shear stress at 1 Pa (10 dyn/cm 2). In order for a single osteocyte process (~ 20 μm long, 50–400 nm in diameter) [10] to receive 10 pN and move a molecular sensor by several nanometers, it is plausible that there is an amplification mechanism
Commentary
such as the one proposed by Weinbaum [11]. Alternatively, local fluctuations from anisotropic and stochastic dynamics may induce a spectrum of stress distributions that might facilitate activation of molecular interactions [12]. Besides the above-mentioned working hypothesis, DHS allows us to consider other mechanisms indirectly linked to mechanical stimulation. Mechanical loading must trigger migration of osteoblasts to the endosteum and periosteum. Because ImP is altered in the intramedullary cavity, it is interesting to ask whether alterations in ImP accelerate migration of osteoblasts and whether any differences exist in the rate of bone formation in the endosteum and periosteum. An additional question is whether bone formation would take place equally throughout the length of the tibia in cortical and trabecular bone or whether any preferential location exists. Lastly, an intriguing question is the importance of local vs. systemic effects in connection to blood circulation and neuronal signaling [7,13]. DHS seems to provide a promising model system for evaluating the role of blood circulation and neuronal signaling in load-driven bone formation. In summary, as the field continues to make advances in our understanding of the physical and structural conditions that stimulate bone formation, safer and perhaps more effective approaches to loadinduced bone formation can be achieved. These developments could not be more timely, as physicians look for alternative means to maintain bone mass in patients once they complete an osteoporosis drug treatment regimen. Several recently-discovered therapies (e.g., Rank-L inhibitors, PTH fragments, and cathepsin K inhibitors) appear to have little to no residual effects once treatment is stopped because they are not stored in the bone tissues in active form. Therefore, bone density is likely to return to baseline once these next-generation therapies are stopped. As the more long-lasting bisphosphonate therapies fall out of favor and newly generated compounds rise in popularity among physicians, the need to maintain bone mass after treatment withdrawal will become more urgent. Therefore, a safe, effective mechanical loading modality potentially provides an alternative treatment for maintaining BMD without inducing deleterious side effects. Although DHS has not been tested clinically, it is not difficult to envision its implementation via a portable loading device that may resemble the inflatable cuff of a blood pressure monitoring system. Its extension from diaphysis to the femoral neck and head seems natural, but its application to spine treatment is less evident. It would be interesting to explore DHS' potential capability not only to enhance bone formation beyond the immediate local environment, but also to stimulate fracture healing without direct application of mechanical loads to the fracture site.
827
[4] Zhang P, Turner CH, Yokota H. Joint loading-driven bone formation and signaling pathways predicted from genome-wide expression profiles. Bone 2009;44:989-98. [5] Zhang P, Su M, Liu Y, Hsu A, Yokota H. Knee loading dynamically alters intramedullary pressure in mouse femora. Bone 2007;40:538-43. [6] Kwon RY, Meays DR, Tang WJ, Frangos JA. Microfluidic enhancement of intramedullary pressure increases interstitial fluid flow and inhibits bone loss in hindlimb suspended mice. J Bone Miner Res 2010;25:1798-807. [7] Bergula AP, Huang W, Frangos JA. Femoral vein ligation increases bone mass in the hindlimb suspended rat. Bone 1999;24:171-7. [8] Fritton SP, Weinbaum S. Fluid and solute transport in bone: flow-induced mechanotransduction. Annu Rev Fluid Mech 2009;41:347-74. [9] Sawakami K, Robling AG, Ai M, Pitner ND, Liu D, Warden SJ, et al. The Wnt co-receptor LRP5 is essential for skeletal mechanotransduction but not for the anabolic bone response to parathyroid hormone treatment. J Biol Chem 2006;281: 23698-711. [10] Sugawara Y, Kamioka H, Honjo T, Tezuka K, Takano-Yamamoto T. Three-dimensional reconstruction of chick calvarial osteocytes and their cell processes using confocal microscopy. Bone 2005;36:877-83. [11] Han Y, Cowen SC, Schaffler MB, Weinbaum S. Mechanotransduction and strain amplification in osteocyte cell processes. Proc Natl Acad Sci U S A 2004;101: 16689-94. [12] Knothe Tate ML. Top down and bottom up engineering of bone. J Biomech 2011;44:304-12. [13] Sample SJ, Behan M, Smith L, Oldenhoff WE, Markel MD, Kalscheur VL, et al. Functional adaptation to loading of a single bone is neuronally regulated and involves multiple bones. J Bone Miner Res 2008;23:1372-81.
Hiroki Yokota Biomechanics and Biomaterials Research Center, Indiana University Purdue University Indianapolis, Indianapolis, IN 46202, USA Department of Biomedical Engineering, Indiana University Purdue University Indianapolis, Indianapolis, IN 46202, USA Department of Anatomy and Cell Biology, Indiana University School of Medicine, Indianapolis, IN 46202, USA Corresponding author at: Biomechanics and Biomaterials Research Center, Department of Biomedical Engineering, Indiana University Purdue University Indianapolis, 723 West Michigan Street, Indianapolis, IN 46202, USA. Fax: +1 317 278 5244. E-mail address: [email protected]. Andrés Tovar Biomechanics and Biomaterials Research Center, Indiana University Purdue University Indianapolis, Indianapolis, IN 46202, USA Department of Mechanical Engineering, Indiana University Purdue University Indianapolis, Indianapolis, IN 46202, USA
Acknowledgment The work is in part supported by NIH R01AR052144 (HY) and R01AR053237 (AR). References [1] Hu M, Cheng J, Qin YX. Dynamic hydraulic flow stimulation on mitigation of trabecular bone loss in a rat functional disuse model. Bone 2012;51:819-25. [2] Qin YX, Kaplan T, Saldanha A, Rubin C. Fluid pressure gradients, arising from oscillations in intramedullary pressure, is correlated with the formation of bone and inhibition of intracortical porosity. J Biomech 2003;36:1427-37. [3] Qin YX, Lam H. Intramedullary pressure and matrix strain induced by oscillatory skeletal muscle stimulation and its potential in adaptation. J Biomech 2009;42: 140-5.
Alexander Robling Biomechanics and Biomaterials Research Center, Indiana University Purdue University Indianapolis, Indianapolis, IN 46202, USA Department of Biomedical Engineering, Indiana University Purdue University Indianapolis, Indianapolis, IN 46202, USA Department of Anatomy and Cell Biology, Indiana University School of Medicine, Indianapolis, IN 46202, USA 8 June 2012