516Views & Citations
The need to image and measure the mechanical properties of skin, scar, tumors, extracellular matrices (ECMs) and wound tissue has been a goal of researchers since the 1970s. A variety of methods have been used to image and evaluate the mechanical properties of skin over the last 40 years including uniaxial and biaxial tensile testing, indentation and rotational tests, ultrasound elastography (UE), optical coherence tomography (OCT), optical coherence elastography (OCE), and vibrational analysis combined with OCT.
We recently reported the use of OCT and vibrational analysis to identify the differences between skin and scar tissue. This paper describes the use of vibrational OCT to image and determine the physical characteristics of the margins of a skin lesion. We report that differences in the collagen orientation between skin and scar tissue appear to increase the modulus of the collagen network in scar tissue by a factor of about 2. Using images generated by OCT, and maps of the modulus as a function of position, it is possible to identify the edges of scar tissue and other skin lesions using vibrational OCT.
Keywords: Collagen, Imaging vibrational optical coherence tomography (VOCT), Optical coherence tomography (OCT), Mechanical properties, Decellularized dermis, Skin, Scar, Scar margins, Modulus, Natural frequency, Imaging vibrational optical cohesion tomography Imaging, Vibrational optical coherence tomography (VOCT),
Mechanobiology and physiology play important roles in normal tissue homeostasis as well as in repair, regeneration and disease processes [1-5]. Mechanical forces in extracellular matrix (ECM) not only influence the storage and dissipation of mechanical energy during locomotion and blood flow [6-9], they also influence gene expression and tissue remodeling . Extracellular matrices (ECMs) found in musculoskeletal, cardiovascular, dermal and other tissues are under tension during normal physiologic loading, even in the absence of external forces [1,2]. This tension not only fulfills cosmetic functions, but also creates a state of dynamic loading at the collagen fibril-cell interface and at cell-cell attachment points that stimulate conversion of mechanical work (a force moving an object through a distance) into chemical energy (synthesis of high molecular weight cellular components) [1-3].
Changes in the mechanical properties of ECM are known to accompany the onset and progression of several highly prevalent diseases, including atherosclerosis, cirrhosis, and cancer. At least one publication underscores the relationship between cancer and tissue fibrosis . A technique to image and measure of the mechanical properties of tissues may enable early diagnosis of some of these conditions.
Extracellular matrix in the tumor stroma is characterized by remodeling and stiffening; tissue stiffness has been used to detect cancer [12,13]. ECM stiffening has been reported to enhance cell growth and survival, and promote cell migration ; ECM rigidity disrupts tissue morphogenesis by increasing cell tension . Reduction of cell tension has been reported to repress the malignant behavior of mammary epithelial cells (MECs) and normalized the behavior of breast cancer cells in culture . ECM stiffening in tumors may be related to changes in the phosphorelay pathways, and ECM tension has been postulated to drive tumor formation . The need to measure the mechanical properties of skin, scar, tumors, extracellular matrices (ECMs) and wound tissue has been a goal of researchers since the 1970s. The pioneering work of Yamada  and Fung  illustrated how difficult this goal would be since the behavior of human ECM depends on strain-rate, direction of testing and is time-dependent . A variety of methods have been used to evaluate the mechanical properties of skin over the last 40 years including uniaxial and biaxial tensile testing, indentation and rotational tests, ultrasound elastography (UE), optical coherence tomography (OCT), optical coherence elastography (OCE), and vibrational analysis combined with OCT [22-25]. Many of these techniques require the assumptions that the material is linearly elastic, Poisson’s ratio is close to 0.5 and that viscoelasticity does not dramatically affect the resulting properties of the tissue. However, skin is a non-linear material that is viscoelastic and has upward curvature to the stress-strain curve. This fact makes determination of the stiffness (tangent to the stress-strain curve) and other mechanical properties very difficult to quantitatively analyze since the tangent to the stress-strain curve is constantly changing [22,24,25]. However, despite all of these problems, there is a need to be able to characterize the mechanical properties of skin, since this would give clinicians valuable information about pathological changes that occur during disease processes, the stage of diabetic skin ulcers and the efficacy of cosmetic treatments. In this paper, we will discuss the results of a study to image using OCT and determine the mechanical properties of scar tissue and surrounding skin using vibrational analysis. This paper describes the use of vibrational OCT to image and to determine the physical characteristics of the margins of a skin lesion.
Calibration Sample Preparation
Human decellularized dermis was obtained from allograft tissue as described previously [24-26]. Decellularized human dermal samples were tested after immersion in phosphate buffer solution as described elsewhere [24-26]. All samples were tested wet after soaking in phosphate buffer solution at pH 7.4 for at least 30 minutes. Processing and testing steps were conducted at 22oC.
Human skin and normal scar tissue were evaluated in vivo to demonstrate the clinical use of the OCT and vibrational techniques. The scar tissue was the result of a small thermal burn wound smaller than the size of a dime that was depigmented after healing (Figure 1). Mechanical Testing
Incremental Stress-Strain Tensile Measurements in Vitro
Samples were tested in uniaxial tension at 22oC by adding a strain increment and then measuring the load before an additional strain step was added as described previously [23-26]. Varying axial deformations of between 1 and 14% were applied through adjustment of a graduated translation stage. The resulting axial force (F) was measured by the force gage and recorded for subsequent calculations. Stress values were calculated from the force divided by the cross-sectional area. Strains were calculated by dividing the change in length by the original length based on the movement of the translational stage after each strain increment was added. The tensile modulus was calculated from a tangent drawn to the stress-strain curve at the strain increment used.
OCT and Vibrational Analysis in Vitro
Transverse forces were applied to the sample by positioning an acoustic loudspeaker (Intervox S225RA-40) beneath the sample. A function generator (Agilent) was used to drive the speaker with sinusoidal waveforms at varying amplitude and frequency.
Transverse sample displacement was measured by spectral-domain optical coherence tomography (SD-OCT), a non-contact, interferometric technique as discussed previously [25,26]. The SD-OCT system uses a fiber-coupled superluminescent diode light source with 1325 nm center wavelength and 100 nm bandwidth (full-width at half maximum) [25,26].
The resonant frequency of each sample was initially estimated at a signal point by measuring the transverse displacement resulting from sinusoidal driving frequencies ranging from 50 Hz to 1000 Hz, in steps of 50 Hz. Once the region where the maximum frequency was identified, smaller steps of 10 Hz were used to more accurately identify the peak frequency and the actual resonant frequency, fn (Figure 2).
The modulus from in vitro vibrational studies was determined using equation (1) where m, L and A are the sample mass, length and cross-sectional area.
In Vivo Determination of the Resonant Frequency of Skin and Scar Tissue
In vivo studies on the mechanical properties of skin and healed scar tissue were conducted by hard wiring a 24 mm x 14 mm rectangular speaker (Digi-Key, Thief River Falls, MN) to a Samsung cell phone. A frequency generating app was downloaded from the Google Play Store onto the cell phone. This app was capable of driving the speaker between 10 and 20,000 Hz. The speaker was taped to the skin using surgical tape and it was used to generate a sinusoidal sound wave that vibrated the skin. During in vivo measurements, no sensation of the light or sound impinging on the skin was felt. The sound intensity was low enough so that it could not be detected unless the speaker was placed near the subject’s ear to make sure it was energized.
The Digi-Key speaker was used for in vivo measurements in place of the Intervox speaker described for the in vitro studies above. The speaker was located about 2.5 cm from the where the incident light beam contacted the skin and did not interfere with impingement of the light on the skin. The location of the incident beam on the skin influenced the extent of the displacement but not the resonant frequency (data not shown). The optical signal generated by vibrating the skin with the Digi-Key speaker was then processed in the same manner as was done for in vitro studies and the resonant frequency was obtained by determination of the frequency at which the displacement was maximized. Measurements made with the Digi-Key speaker were made on decellularized human dermis in vitro and the resonant frequency determined using this speaker was similar to that measured with the Intervox speaker.
A variety of samples made from silicone rubber, decellularized dermis, and chemically modified decellularized dermis were tested in uniaxial tension and using vibrational analysis to establish a calibration curve between the moduli calculated from tensile measurements and those derived from vibrational measurements in vitro. These results have been published elsewhere [25,26].
The relationship between the modulus measured using vibrational and tensile measurements was reported to be approximately linear and the equation of the line was found to be:
Ev=1.026 Et + 0.0046 (2)
where, Ev and Et are the moduli measured using vibrational and tensile measurements, respectively and are in MPas. The correlation coefficient between these moduli is 0.984 as previously reported [25,26]. The relationship between tensile and vibrational moduli was approximated using equation (2). The material behavior was reported to be reversible for strains less than about 14% for up to three cycles of tensile testing [25,26].
Photographic and OCT images of skin and scar were made using a Samsung cell phone, and a Lumedica OQ labscope1.0 (Lumedica, Inc., Durham, NC) operating in the scanning mode, respectively.
Vibrational OCT was used to measure the mechanical properties of human skin and scar in vivo. The results of previously published data on decellularized dermis and silicone rubber were used to calibrate the in vitro and in vivo studies. The resonant frequency was measured from vibrational OCT studies as shown in Figure 2. Resonant frequencies were obtained by determining the frequency at which the maximum displacement was observed by measurements at a single point (Figure 2). For decellularized dermis, the value of the resonant frequency at 5% strain was calculated to be 150 Hz as reported previously [25,26]. The modulus of decellularized dermis at 5% strain was calculated from equation (1) as previously described and compared to the tensile modulus using equation (2) and the calibration curve (Figure 3) as described previously [25,26].
Photographic and OCT images of both skin and scar are shown in Figures 1 and 4, respectively. Photographic images showed that the scar is approximately 7 mm in diameter and was clearly demarcated from the surrounding skin by differences in pigmentation. The edges of the scar were marked N, S, E and W to determine whether differences in the mechanical properties could be measured in comparison to the normal skin (skin) and the scar tissue proper (scar). OCT scanning images of the skin and scar cross-sections are shown in Figure 4. The normal skin appeared to have surface hills and valleys while the scar tissue appeared smoother in texture of the surface.
Plots of weighted displacement versus frequency for normal skin, scar, and one of the edges of the scar are shown in Figure 5. The plot of weighted displacement versus frequency for scar tissue is shifted to the right as compared to that of normal skin (higher resonant frequency). The resonant frequency of normal skin (Skin) was found to be 100 Hz while that of scar tissue (Scar) was 220 Hz. The edges of the scar (N, S, E, W) had resonant frequencies of both normal skin (90-100 Hz) and scar tissue (220-230 Hz) as shown in Figure 5 and in Table 1.
The resonant frequency for decellularized dermis was found to be dependent on sample thickness. However, the product of E, the modulus, and the sample thickness, d, was found to be related to the resonant frequency as shown in Figure 6. The resonant frequency and thickness in vivo were determined from vibrational OCT and the modulus was calculated from Figure 6. The calculated modulus of normal skin was determined to be 2.98MPa and that of scar tissue was 6.85 MPa. These values are close to the values reported for decellularized dermis at 5% strain (skin) and 14% strain (scar) reported previously [25,26]. The edges of the scar had moduli that were similar to both normal skin and scar tissue.
The inability of researchers and clinicians to understand how to correlate changes in images of skin lesions with the content and organization of the ECM hampers the development of new non-invasive methods to understand and diagnose the pathogenesis and nature of skin disorders. While the mechanical properties of skin are complex, much progress has been made in understanding the strain-rate dependence, non-linearity and compressibility of this tissue [25,26].
We recently reported the use of OCT and vibrational analysis to identify the differences between skin and scar tissue [25-27]. The correlation between modulus measurements on decellularized human dermis made using standard tensile testing in vitro and vibrational OCT suggest that measurements made using vibrational OCT give results that are consistent with tensile testing, a “gold standard method” for measuring mechanical properties of skin. Without comparison to a standard technique, moduli measurements made with new methods such as vibrational OCT cannot be validated.
Figure 7. Photographic image and modulus values for normal human skin and scar. Photographic image for skin and scar are shown on the left. Modulus values for normal skin (2.98 MPa) and scar (6.85 MPa) are shown on the right as well as for the points N,E, S and W. The edges of the scar have two resonant frequencies and modulus values corresponding to that of normal skin and scar were determined from Figure 6 and are listed in Table 1.
Tensile incremental and constant rate-of-strain measurements made on tissues have been the gold standard for determination of the mechanical properties of tissues for decades [20-22]. Many techniques require the assumption that the tissue density is near 1.0 and that Poisson’s ratio is 0.5. The later has been shown to vary between 0.38 and 0.75 for decellularized dermis [25-28] and the tissue density for human skin is about 1.2 g/cc. The assumption that Poisson’s ratio is 0.5 will lead to errors in modulus calculations. However, by measuring the resonant frequency using vibrational analysis and OCT and using equations (1) and (2), the tensile modulus can be determined for skin without the need to make any assumptions. In addition, the results of vibrational studies on pig skin indicate that properties of both the elastic and collagen fiber networks can be measured in the same experiment, which constitute the major contributors to mechanical behavior of skin [1,2,29].
By calibrating our in vivo system for measuring the resonant frequency of normal and scar tissue, we are able to show that the resonant frequency of scar tissue is 220-230 Hz, an increased value from than that found for normal skin (90-100 Hz). The resulting increase in modulus for scar tissue compared to normal skin is consistent with a previous report that suggests that hypertrophic scar tissue appears stiffer than normal skin due to its inability to deform during mechanical loading . The fact that the edges of the scar tissue show resonant frequencies of both normal skin and scar tissue suggest that two different collagen networks exist at the edges of the scar tissue. Normal skin contains an almost biaxial network of collagen fibers that are able to orient under increasing loads with the direction of tension [1,2]. The modulus of the collagen network in normal skin increases with increased deformation until the network becomes fully recruited with the loading direction. In contrast, the modulus of scar tissue is much higher at low strains since the collagen fibers in scar tissue are more disorganized and have difficulty reorienting during tensile deformation. In is likely that the edges of the scar have a combination of normal skin with biaxial orientation and disorganized collagen in the scar that give rise to both the normal skin modulus as well as the higher scar modulus (see Figure 7).
The ability to measure differences in the collagen orientation suggests that vibrational OCT may be useful in imaging and measuring mechanical differences that occur at the interface between normal skin and skin lesions. This may be useful in characterizing the margins of benign and malignant skin lesions as well as the extent of healing of diabetic skin ulcers and other wounds. The technique may also be useful in evaluating the efficacy of treatments to skin and tumors.
Using vibrational OCT the resonant frequency and moduli of collagen in skin and scar tissue can be measured non-invasively and non-destructively. The numbers generated reflect to a first approximation the elastic moduli and do not depend on measurement of other parameters. Differences in the collagen orientation between skin and scar appear to alter the modulus of the collagen network by a factor of about 2. The technique in vitro is calibrated using incremental tensile measurements and vibrational OCT results on decellularized human dermis. Using images generated by OCT, and maps of the modulus as a function of position, it may be possible to determine the margins of scars, tumors, as well as to evaluate the effects of cosmetic treatments to the skin.
- Silver FH, Siperko LM, Seehra GP (2002) Mechanobiology of force transduction in dermis. Skin Res Technol 8: 1-21.
- Silver FH, DeVore D, Siperko LM (2003) Invited Review: Role of mechanophysiology in aging of ECM: effects of changes in mechanochemical transduction. J Appl Physiol 95: 2134-2141.
- Silver FH (2006) Mechanosensing and Mechanochemical Transduction in Extracellular Matrix, Springer, NY.
- Silver FH, Siperko LM (2003) Mechanosensing and Mechanochemical Transduction. Crit Rev Biomed Eng 31: 255-331.
- Moore SW (2003) Scrambled eggs: mechanical forces as ecological factors in early development. Evolut Develop 5: 61-66.
- Silver FH, Snowhill PB, Foran D (2003) Mechanical behavior of vessel wall: A comparative study of aorta, vena cava, and carotid artery. Ann Biomed Eng 31: 793-803.
- Freeman JW, Silver FH (2004) Elastic energy storage in unmineralized and mineralized extracellular matrices (ECMs): A comparison between molecular modeling and experimental measurements. J Theor Biol 229: 371-381.
- Horvath I, Foran DJ, Silver FH (2005) Energy Analysis of Flow Induced Harmonic Motion in Blood Vessel Walls, Cardiovasc Eng 5: 21-28.
- Freeman JW, Silver FH (2004) Elastic energy storage in unmineralized and mineralized extracellular matrices (ECMs): A comparison between molecular modeling and experimental measurements. J Theor Biol 229: 371-381.
- Chiquet M (1999) Regulation of an extracellular matrix gene expression by mechanical stress. Matrix Biol 18: 417-426.
- Leventhal KR, Yu H, Kass L, Latkins JN, Egeblad M, Erler JT, Fong SFT, Csiszar K, Giacci A, Weninger W, Yamauchi M, Gassar DL, Weaver, VM (2009) Matrix crosslinking forces tumor progression by enhancing integrin signaling. Cell 139: 891-906.
- Butcher DT, Allston T, Weaver VM (2009) A tense situation: forcing tumor progression. Nat. Rev. Cancer 9: 108-122.
- Sinkus R, Lorenzen J, Shrader D, Lorenzen M, Dargatz M, Holtz D (2000) High-resolution tensor MR elastography for breast tumour detection. Phys Med Biol 45: 1649-1664.
- Lo GM, Wang HB, Dembo M, Wang YL (2000) Cell movement is guided by the rigidity of the substrate. Biophys J 79: 144-152.
- Paszek MJ, Zahir N, Johnson KR, Lakins JN, Rosenberg GI, et al. (2005) Tensional homeostasis and the malignant phenotype. Cancer Cell 8: 241-254.
- Snowhill PB, Foran DJ, Silver, FH (2004) A mechanical model of porcine vascular tissues-Part I. Determination of macromolecular component arrangement and volume fractions. Cardiovasc Eng 4: 281-294.
- Snowhill PB, Silver FH (2005) A mechanical model of porcine vascular tissues-Part II: Stress-strain and mechanical properties of juvenile porcine blood vessels. Cardiovasc Eng 5: 157-169.
- Jodele S, Blavier L, Yoon JM, Declerck, YA (2006) Modifying the soil to affect the seed: role of stromal-derived matrix metaloproteinases in cancer progression. Cancer Metastasis Rev 25: 35-43.
- Ramaswamy, S, Ross, KN, Lander, ES, and Golub, TR (2003) A molecular signature of metastasis in primary solid tumors. Nat Genet 33: 49-54.
- Yamada H (1970) Strength of Biological Materials, Williams and Wilkins, Baltimore, MD.
- Fung YC (1993) Biomechanics: Mechanical Properties of Living Tissue, Second Edition, Springer, NY.
- Dunn MG, Silver FH (1983) Viscoelastic behavior of human connective tissues: Relative contribution of viscous and elastic components. Conn Tis Res 12: 59-70.
- Silver FH, Shah R (2016) Measurement of mechanical properties of natural and engineered implants. Adv Tissue Eng Reg Med 1: 1-9.
- Shah R, DeVore D, Pierce MG (2016) Morphomechanics of dermis-A method for non-destructive testing of collagenous tissues. Skin Res Tech.
- Shah R, Pierce MC, Silver FH (2017) A method for non-destructive mechanical testing of tissues and implants. J Biomed Mat Res 105A: 5-22.
- Silver FH, Freeman J, DeVore D (2001) Viscoelastic properties of human skin and processed dermis. Skin Res Tech 7: 18-23.
- Silver FH, Silver (2017) Non-invasive viscoelastic behavior of human skin and decellularized dermis using vibrational OCT. Derm Clin Res 3: 174-179.
- Dunn MG, Silver FH, Swann DA (1985) Mechanical analysis of hypertrophic scar tissue: Structural basis for apparent increased rigidity. J Invest Dermatol 84: 9-13.
- Shah RG, Silver FH (2017) Vibrational analysis of extracellular matrix scaffolds: comparison of skin, dermis, cartilage and subchondral bone. Adv Tiss Eng Regen Med 2: 00048.
- Oncology Clinics and Research (ISSN: 2643-055X)
- Journal of Cell Signaling & Damage-Associated Molecular Patterns
- International Journal of Clinical Case Studies and Reports (ISSN:2641-5771)
- Journal of Cardiology and Diagnostics Research (ISSN:2639-4634)
- Journal of Alcoholism Clinical Research
- International Journal of Anaesthesia and Research (ISSN:2641-399X)
- International Journal of Surgery and Invasive Procedures (ISSN:2640-0820)