Trachea Mechanics for Tissue Engineering Design - ACS Publications

Review Cite This: ACS Biomater. Sci. Eng. XXXX, XXX, XXX−XXX

Trachea Mechanics for Tissue Engineering Design Elizabeth M. Boazak† and Debra T. Auguste*,†,‡ †

Department of Biomedical Engineering, The City College of New York, Steinman Hall, 160 Convent Avenue, New York, New York 10031, United States ‡ Department of Chemical Engineering, Northeastern University, 360 Huntington Avenue, Boston, Massachusetts 02115, United States ABSTRACT: Trachea replacement for nonoperable defects remains an unsolved problem due to complications with stenosis and mechanical insufficiency. While native trachea has anisotropic mechanical properties, the vast majority of engineered constructs focus on uniform cartilaginous-like conduits. These conduits often lack quantitative mechanical analysis at the construct level, which limits analysis of functional outcomes in vivo, as well as comparisons across studies. This review aims to present a clear picture of native tracheal mechanics at the tissue and organ level, as well as loading conditions to establish design criteria for trachea replacements. We further explore the implications of failing to match native properties with regards to implant collapse, stenosis, and infection. KEYWORDS: trachea mechanics, cartilage rings, annular ligament, trachealis muscle, tracheal compliance, engineered trachea, structural anisotropy materials showing inadequate circumferential strength.14 However, even in cases where mechanical strength may be sufficient, stenosis is a serious problem.7,8 The majority of reviews reference early work15 pointing to the importance of the anisotropic mechanical properties of the native trachea. In 1929, physiologists first suggested that a completely rigid tube would not allow for normal lung ventilation.16 Yet, despite the repeated urging that longitudinal extensibility and radial rigidity will be required in a trachea replacement, there has been a scarcity of quantitative mechanical data reported for engineered tracheae and limited discussion of the design criteria for a functional implant. The goal of this review is to present a clear picture of native trachea mechanical properties and loading conditions, and to explore the implications of failing to match these properties with regards to implant collapse, stenosis, and infection.

1. INTRODUCTION The need for a trachea replacement can arise from a variety of sources: neoplastic lesions (primarily adenoid cystic carcinoma1), tracheomalacia, long segment tracheal stenosis, and traumatic injury. Tracheal repair is also required for congenital abnormalities, which include tumors, bronchotracheal and trachea-esophageal fistula, tracheal atresia, and tracheal agenesis.2 Cervical and thoracic tumors often result in airway damage, either intrinsically or as an effect of emergency intubations and attempts to resect the tumor.3,4 Surgical repair is not feasible for defects comprising over 50% of the adult tracheal length or 30% of the length in infants and small children, due to excessive anastomotic tension.1,5,6 Trachea replacement for nonoperable defects remains an unsolved problem due to complications with the numerous prostheses developed, such as stenosis, implant weakening, and inconsistent functional outcomes, even when complete epithelialization is observed.7,8 Increasingly, the focus has turned to tissue engineering for a solution. Donor tissue for transplant or decellularization is limited, and immunosuppression is particularly undesirable, as a large number of patients require a replacement due to carcinoma.1 Solid synthetic tubes are not incorporated well and typically dislodge, migrate, or become occluded.1 Several reviews9−13 have chronicled the trends in tracheal replacement in the past 10 years (2006−2016), and have collectively emphasized the need for epithelialization, vascularization (particularly to support the epithelium), and addressing stenosis. Tracheal collapse is recognized as a main mode of failure for trachea implants, with nearly all investigated © XXXX American Chemical Society

2. TRACHEA STRUCTURE AND TISSUE MECHANICAL PROPERTIES The adult human trachea is on average 1.8 cm in diameter and 12 cm in length, although it lengthens and widens during inhalation and narrows and shortens during exhalation.17,18 Neonatal diameters and lengths are significantly smaller, ranging from 3 to 4.5 mm19 and 20 to 30 mm,20 respectively. In the developing child, the trachea is initially slightly cone Received: October 25, 2017 Accepted: February 27, 2018 Published: February 27, 2018 A

DOI: 10.1021/acsbiomaterials.7b00738 ACS Biomater. Sci. Eng. XXXX, XXX, XXX−XXX

Review

ACS Biomaterials Science & Engineering

yet it is the unique combination and organization of a variety of proteoglycans that engenders cartilage compressive properties;25 members of all classes of small leucine-rich proteoglycans (SLRPs) are found in cartilage, including decorin,26 biglycan,27 fibromodulin,28 lumican,29 epiphycan,30 and perlecan.31 Cartilage is typically modeled as a biphasic material to account for the behavior and interaction of the solid matrix phase and the interstitial fluid phase.32 Its highly negatively charged proteoglycan content enables fluid retention and regulates permeability of the tissue.33 Cartilage is comprised of 60−85% water and dissolved electrolytes.34 Cartilage mechanical function is governed by the matrix, rather than cellular, properties. The interaction between water and the proteoglycans generates drag, which opposes the flow of water out of the tissue under compression.32 The aggregate elastic modulus is the modulus when the fluid flow out of the tissue has reached equilibrium. Thus, the aggregate or equilibrium modulus is then testing mainly the tensile or compressive properties of the collagen network, rather than the complex tissue response to mechanical loading. Mechanical tests performed at slow loading rates are attempting to capture the elastic properties of the material. In vivo, tracheal rings are exposed to both sustained and transient loading. Contraction of the trachealis muscle can apply sustained loads, while coughing or other situations resulting in maximum expiratory flow result in higher magnitude transient loads.35 The effects of transient loads on cartilage deformation depend on its viscoelastic properties. The cartilage rings are responsible for holding the trachea open in spite of changes in the interthoracic pressure which occur during respiration.36 Yet, the rings still must be capable of significant deformations to allow for changes in cross-sectional area. It is believed that changes in overall stiffness of the trachea may be a contributing factor to the altered lung function seen in obstructive lung diseases.37 Adult human trachea rings are reported to have a tensile equilibrium modulus between 1 and 20 MPa.35,38,39 This large range is due to both variations in measurement methods and interindividual variation. The tensile modulus of human tracheal cartilage increases with age, while hydroxyproline content, reflecting collagen content, decreases, and proteoglycan content remains stable.35,40 Table 1 contains values reported in the literature for mechanical properties of tracheal cartilage. Both tensile and compressive properties are relevant to the performance of tracheal cartilage, as rings are loaded in such a way that bending generates tensile loads on the outer regions of the rings and compressive loads on the inner regions (Figure 2).35 2.2. Trachealis Muscle. The trachealis muscle is the active component of the trachea, capable of responding to external stimuli.21 It is a smooth muscle, responsible for changing trachea diameter in the cough reflex or cases such as the asthmatic response. The trachealis muscle contains stretch receptors which respond to transverse strain, but not longitudinal.41 Muscle contraction helps prevent compression and enhance maximum airflow during deep and forced breathing.42 Smooth muscle cells contribute to both the active and the passive mechanical properties of the trachealis muscle, as it is a highly cellular tissue. Trachealis matrix synthesis and muscle contraction has been primarily studied in the context of asthma. Generally matrix synthesis is increased in asthmatic patients. Specifically, collagen I, III, and V, fibronectin, tenascin, hyaluronan, versican, perlecan, and laminin α2/β2 synthesis is increased, while collagen type IV and elastin synthesis is

shaped, with the rings near the top having a larger circumference than those near the carina. In neonates the difference in circumference is approximately 1 mm, while in older children, this difference is only 0.1 mm.18 Both the length of the cartilage rings and the length of the muscle increase with age, such that the ratio of cartilage to muscle length about the circumference remains constant at a ratio of approximately 5:1.18 Wailoo and Emery describe three phases of tracheal growth, in which the rate of increase in tracheal internal circumference coincides with the increase in crown-rump length.18 During phase 1, corresponding to the final 12 weeks of gestation, the internal tracheal circumference enlarges by 4 mm, while the crown-rump length increases by 15 cm. Thus, phase 1 is characterized by a 2.6 mm circumference increase per cm increase in crown-rump length. In phase 2, the growth rate is nearly doubled at 5 mm per 10 cm increase in crown-rump length. Phase 3, occurring after puberty, again is marked by the slower growth that characterized phase 1. Horseshoe-shaped cartilaginous rings, annular ligament, trachealis muscle, and epithelium are the primary structural components of the trachea (Figure 1).17,21 While the

Figure 1. The trachea is comprised of horseshoe shaped cartilaginous rings connected longitudinally by annular ligament and bridged by smooth muscle. The lumen is lined with ciliated epithelium.

epithelium plays an essential physiological role in preventing infection, it does not contribute structurally to maintaining a patent airway. Tracheal cartilage, smooth muscle, and annular ligament are all highly anisotropic structural components that engender the longitudinal extensibility and lateral rigidity of the native trachea.15,21 The rings, of which there are 18−22,22 are responsible for holding the trachea open in spite of the changes in interthoracic pressure which occur during respiration, while the muscle and connective ligaments allow for changes in diameter and length.17 2.1. Tracheal Cartilage. The tracheal rings, composed of hyaline cartilage, comprise up to five-sixths of the trachea circumference. Wachsmuth et al. compared various types of human cartilage23 and found that tracheal cartilage has high cellularity, as compared to articular hyaline cartilage. Collagens types II, III, V, VI, and X are present in tracheal cartilage, with some variation between the ring margins and center. For histological images specifically comparing tracheal cartilage to other types of hyaline cartilage, we recommend referring to the Wachsmuth publication. Collagen type II is the predominate collagen in hyaline cartilage, while type I, present in fibrocartilage and most engineered cartilage, is notably absent.24 Tracheal cartilage exhibits particularly strong and uniform staining for Safrinin O, indicating high proteoglycan content.23 The predominant proteoglycan in hyaline cartilage is aggregan, B


linear elastic curved beam analytical model to estimate Young’s modulus

ring bending with preconditioning

uniaxial tensile test (15−25% strain)

uniaxial tensile test

uniaxial tensile test

ring bending with preconditioning

lamb, n = 6, age 2 months

human, n = 10, age 17−81 years

human, n = 10, age 17−81 years human, n = 4, age 36, 44, 56, 74

Kojima et al. 200353

Roberts et al. 199839

Rains et al. 199235 Lambert et al. 199138

linear elastic

linear aggregate and elastic elastic, neo-Hookean analytical using curved beam theory Fung-type strain energy density function, tensile linear model linear elastic

pig, n = 3

pig, n = 12

Teng et al. 200837

model type linear elastic

cyclical uniaxial tensile test with preconditioning radial tension and compression with preconditioning

test type

human trachea from autopsy, n = 3, age 46, 79, 82 lamb, n = 8 3−6 months

source

uniaxial tensile with preconditioning confined compression, tensile

Safshekan et al. 201622 Hoffman et al. 201651 Trabelsi et al. 201052 Karkhanis 2010

human, n = 30

author

Young’s modulus = 2.5−7.7 MPa

equilibrium modulus 1−15 MPa

equilibrium modulus = 13.6 ± 1.5 MPa (ablumenal superficial zone) and 4.6 ± 1.7 MPa (middle zone)

Young’s modulus = 10.6 ± 1.8 MPa

C1 = 41.9 MPa (main stiffness parameter for linear tensile test). For nonlinear parameters, see reference.37

Despite similar biochemical composition, tissue engineered trachea exhibited much poorer mechanical properties as compared to native tissue. Stress−strain curves were linear for strains up to 10%. Equilibrium tensile modulus decreases with depth. Specimens ranged from 1 to 20 MPa. Stiffness increased significantly with age. Water and hydroxyproline content decrease with age. Calculation methods are expected to underestimate true modulus values. Tensile tests yielded Young’s modulus values from 0.5 to 18 MPa

High interanimal variations. Constant, physiological deformation rate may have resulted in higher values than other studies reporting true equilibrium moduli Higher strength in compression than in tension. Linear stress−strain relation underestimates the stability of trachea by exaggerating the displacement in compression. Tracheal cartilage stiffness may be species specific.

Modulus = 33.50 ± 30.94 MPa (tension) and 17.42 ± 12.24 MPa (compression)

Compressive aggregate modulus = 1.36 ± 0.49 MPa; tensile elastic modulus = 5.62 ± 2.01 MPa Young’s modulus = 3.33 MPa; Poisson ratio = 0.49

key findings elastic moduli increased with age 13.30 ± 5.72 MPa and 20.71 ± 10.17 MPa for young and old samples. Circumferential modulus was significantly different from the longitudinal composite modulus. A linear model may be suitable for cartilage in tension.

reported constants average Young’s modulus = 16.92 ± 8.76 MPa; range of 5.05−39.63 MPa

Table 1. Tracheal Cartilage Mechanical Properties Listed by Date Reported

ACS Biomaterials Science & Engineering Review

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Stiffness increases with age but changes are not significant. Muscle contraction increases stiffness. elastic modulus approximately 15 kPa

The elastic modulus of the trachealis muscle decreased after extended mechanical ventilation. elastic modulus = 195 ± 41 kPa neonatal lamb, in vivo, n = 14, age 1−4 days

guinea pig, n = 5, age 3 months

Miller et al. 200754

Wang et al. 200547

uniaxial quasi-static tensile tests on two orthogonal families of collagen fibers with preconditioning changes in airway diameter were produced using mechanical ventilation uniaxial tensile with preconditioning human trachea from autopsy, n = 3, age 46, 79, 82 years Trabelsi et al. 201052

human, n = 3, age 22, 55, 72 years

reinforced hyperelastic, Holzapfel strain-energy density function calculated with thin walled cylinder assumption linear elastic

notable findings

While older samples were observed to be stiffer, the difference was not statistically significant. Trachealis muscle is more extensible in the circumferential than longitudinal direction. Adventitial and submucosal membranes were less extensible than the trachealis muscle. Tracheal muscle is stiffer in the longitudinal direction.

reported constants

human, n = 30

source

approach

model type D

author

Table 2. Trachealis Muscle Passive Mechanical Properties

3. FORCES ON THE TRACHEA DURING RESPIRATION The forces affecting the airway are important in determining adequate parameters for the evaluation and comparison of engineered tracheae both across studies and with native tissues. The mechanical forces on the trachea are numerous and complex, making precise modeling of the airway very

Safshekan et al. 201622 Teng et al. 201221

decreased.43 While the extracellular matrix may not define the mechanical properties of healthy trachealis muscle, abnormal matrix production could affect tissue mechanics. Stephens et al. studied the force−velocity characteristics in dogs and found that trachealis muscle generates higher active tension than other smooth muscles, but has one of the slowest shortening velocities.44 Panitch et al. investigated the relationship between airway smooth muscle mechanics and aging.45 A few studies have reported passive (the absence of muscle contraction) mechanical properties (Table 2) that may provide insight for scaffold design. The trachealis muscle has highly nonlinear mechanical properties and is typically modeled as a hyperelastic material. Multiple studies confirm that the trachealis muscle is stiffer in the longitudinal direction than the circumferential direction. The anisotropy ratio is further modulated by the degree of shortening present in contracting muscle.46 While multiple studies have observed a stiffening of the trachealis muscle with age, none have achieved statistically significant results.21,22,47 The ability of the trachealis muscle to deform, as compared to the trachea rings, contributes significantly to tracheal compliance, which is discussed at length in Section 4, and also permits tracheal extension. 2.3. Tracheal (Annular) Ligament. Annular ligament connects the cartilage rings and facilitates longitudinal extensibility. Ligaments are comprised primarily of type I collagen, although have higher type III content than tendons.48 Ligaments also tend to be more metabolically active than tendons, and have higher glycosaminoglycan content.49 While differences in biochemical composition between individual ligaments have been observed,49 we are unaware of any studies directly assessing the matrix components of tracheal ligament. Measuring the longitudinal mechanical properties of annular ligament is challenging due to the short inter-ring length. In modeling, the properties of trachealis muscle are used as an approximation for trachea annular ligaments.50 Recently, Safshekan et al. evaluated annular ligament by testing whole strips of tracheal tissue and mathematically accounting for the hard cartilage segments.22 While they observed samples from older donors were stiffer, this difference was not statistically significant. Longitudinal behavior of composite trachea samples containing both ligament and cartilage rings was further investigated, and is discussed in section 5.

uniaxial tensile with preconditioning (longitudinal) uniaxial tensile

Figure 2. Tracheal ring bending results in the generation of complex loads within the cartilage. Interior regions are subject to compression, while exterior regions are loaded in tension.

isotropic, hyperelastic tangent moduli at varying stretch ratios

Yeoh: a1 = 0.063, a2 = 0.394, a3 = −0.171; Mooney−Rivlin: a1 = 1.164, a2 = −1.223 for stretch ratios of 1, 1.05, 1.1, 1.15, and 1.2: longitudinal tangent modulus = 5.7, 9.6, 16.3, 631.8 kPa and 34 364 kPa; circumferential tangent modulus = 4.6, 6.7, 10.9, 15.7, and 32.2 kPa C10 = 0.877 kPa, k1 = 0.154 kPa and k4 = 34.157 (longitudinal), k3 = 0.347 kPa, and k4 = 34.157 (transversal)



Review


from the EPP. Mead et al. were pioneers in quantifying the physical parameters affecting the trachea.56,57 Key aspects of this work pertaining to tracheal compression were later summarized in the work of Macklem et al.42 It is difficult to identify an individual location prone to collapse, as any given point experiences a range of applied transmural pressures.58 The extratracheal pressure in the thoracic cavity is usually negative, similar to the pleural pressure. The pleural cavity is the fluid filled space between the lungs and the pleural membrane, which is attached to the chest wall. Transpulmonary pressures during quiet breathing vary between 5 and 8 cm H2O.59 However, in the case of forced expiration, the pleural pressure can become positive. Muscle strength determines the maximum pleural pressure that an individual can achieve. The mean maximal inspiratory (MIP) and expiratory (MEP) pressures are 107 ± 26 and 114 ± 35 cm H2O for men, and 76 ± 25 and 86 ± 22 cm H2O for women, with no significant differences between adolescents and adults.60 In full term infants, MIP and MEP have been measured at 70.3 ± 18.7 cm H2O and 52.9 ± 13.4 cm H2O, respectively.61 The extratracheal pressure in the neck is approximately atmospheric, and the pressures at the cervicothoracic junction differ significantly from both compartments. To further complicate the system, the trachea slides considerably between compartments during respiration. These anatomical shifts are on the same time scale as changes in intratracheal pressure, and may therefore be responsible for the changes in transmural pressure experienced at any given tracheal location.58 The changes in normal trachea diameter are barely measurable during quiet respiration. However, in abnormal situations, collapse can occur during both inspiration and expiration, with expiratory collapse more common.58 Wittenborg et al. proposed that collapse is driven by increased transmural pressures as a result of increased respiratory effort in the face of abnormal airway resistance. Airway resistance is equal to the pressure drop driving the airflow divided by the volumetric flow rate. Mead et al. showed that relatively large increases in tracheal resistance occur during expiration as compared to inspiration, which aligns with more prevalent expiratory collapse in infants with respiratory abnormalities.57,58 This difference in airway resistance between inspiration and expiration is additionally supported by bronchoscopic observations of reduced tracheal cross section during rapid expiration and partial trachea collapse during coughing.62 Inspiratory collapse is caused by the generation of more negative intratracheal pressures than normal, while expiratory collapse is caused by the aforementioned increase in respiratory effort and pleural pressure, as shown in Figure 4. Considering that increased resistance at any location in the trachea can lead to the development of abnormally high transmural pressures, fibrotic stenosis has a bearing on the success of an implant beyond the intrinsic reduction in diameter. A construct with appropriate mechanical properties could undergo collapse due to abnormal loading conditions. The relationship between flow rate Q, airway resistance R, and transmural pressure P, such that P = QR, means that the respiratory rate affects both the transmural pressures developed and the time over which loads are normally applied. The average adult takes 12−15 breaths per minute (4.4 s/breath), while the average newborn has a respiration rate of 38 breaths per minute (1.6 s/breath) during quiet breathing. Healthy young adult respiratory rates increase to around 35−45 breaths

challenging. Physical properties vary with muscle tone; muscle contraction can increase airway resistance during quiet breathing but decrease resistance by reducing compliance during forced expiration or cough.55 During respiration, air is driven in and out of the lungs by pressure gradients (Figure 3). In inspiration, the diaphragm

Figure 3. Mechanics of inspiration and expiration. Air is driven into and out of the lungs when diaphragm contraction and relaxation changes the volume of the chest cavity.

contracts, decreasing pressure in the thoracic cavity. This expands the lungs, and hence decreases the pressure within the lungs. Hence, two gradients are formed: one from the lung to the mouth, and another across the airway walls (Figure 4).

Figure 4. Normal and collapsing trachea during inspiration and expiration. Reproduced with permission from reference 58. Copyright 1967 Radiological Society of North America.

During expiration, the diaphragm relaxes, increasing the pressure in the lungs, and driving air out. When maximal expiratory flow is reached, an equal pressure point (EPP) develops. Airway compression does not directly correspond to the location of the EPP; invagination of the posterior membrane and initiation of compression occurs downstream E


Review


Figure 5. Compilation of change in volume vs pressure curves for trachea reported in the literature. Tracheal compliance, the slope of the pressure− volume curve, is nonlinear. Curves reflect both inflating and deflating compliance at positive and negative pressures, respectively. The change in volume, shown on the y-axis, is normalized by initial volume to facilitate comparison across studies. Data points were approximated from plots and publications, and normalized from information given in the methods where necessary and possible.

per minute, although those of elite athletes can be higher.63 Under stressed conditions, infant respiratory rates can increase to 61 breaths per minute, along with some increase in tidal volume, although maximum respiratory rates cannot be maintained for long time periods.36,64 A thorough review by Fleming et al. presents median respiratory rates for children age 0−18 years.65 Tracheal rings, or constructs, must recoil from the applied loads on an appropriate time scale. Testing of engineered constructs at physiologically relevant loading rates under cyclic conditions may be important for the validation of engineered tracheal constructs, particularly at rates representing stressed conditions, to determine if adequate recoil is possible in the case of deformation.66

in the airway leads to respiratory discomfort, airway blockage, and predisposition for infection.71,72 The trachealis muscle contributes most significantly to the changes in cross-sectional area: 91% in dogs, 93% in cats, and 91% in rabbits, while the remainder is contributed by the bending of the cartilage rings.73,74 Trachealis muscle contraction helps regulate changes in cross-sectional area. Fluid−solid interface modeling indicates that stents typically prevent tracheal muscle deflections and alter flow patterns at the top of the stent in ways that may lead to mucous accumulation.72 The insertion of an implant unable to change cross-sectional profile along with the native tissue would also likely establish undesirable flow patterns. In addition to promoting mucous accumulation, excessive radial rigidity could result in radial mechanical mismatch at the attachment sites. From the end of forced inspiration to the end of forced expiration in adults, 32% and 13% mean reductions in anteroposterior and transverse diameter, respectively, were observed.75 Values for individuals ranged from 11 to 61%. In healthy children, 20−50% variations in diameter have been observed during crying or struggling.58 While not a direct physiological loading scenario, radial compression studies may be relevant in the validation of constructs engineered to allow for radial deformation. Force to achieve relevant reductions in diameter as compared to native tissues as well as the ability to recoil should be considered. Compliance testing evaluates the mechanical behavior of the trachea in a more physiologically relevant loading scenario than mechanical testing of individual tissue components or radial compression. Compliance accounts for the contributions of both the rings and muscle to the radial properties at a whole organ or construct level. Airway compliance is defined as the change in airway volume over the change in transmural pressure, (ΔVA/ΔPtm, the slope of a volume pressure curve).55 Volume pressure curves frequently report a ΔVA normalized by the original volume to better enable comparisons across

4. WHOLE ORGAN MECHANICAL PROPERTIES AND IMPLICATIONS OF IMPLANT MECHANICAL MISMATCH Achieving closely matched, anisotropic mechanical properties in a trachea replacement will likely be important for graft patency and functionality. If the implant stiffness is mismatched, granulation at the sites of anastomosis may induce fatal obstruction or bleeding.67 As there are significant loads and deformations both longitudinally and radially in normal tracheal function, mechanical mismatch in either aspect could contribute to excess scarring. Mechanical forces are known to play an important role in fibrosis, through T-cell mediated prolonged inflammation.68 Additionally, tracheal compliance and extensibility are important parameters in normal physiology, affecting respiration and airway regulation. The following subsections explore the implications of failing to account for the native radial and longitudinal properties. 4.1. Radial Geometry and Compliance. The ability of the trachea to resist collapse is balanced by the requirement for deformation under appropriate transmural pressures in order to sufficiently accelerate air and clear mucous.69,70 Excess mucous F


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loading. Uniform cartilaginous tubes will likely struggle to achieve long-term structural integrity even with ideal, initial scaffold material properties and culture conditions. Long-term success of tissue engineered products will depend on the challenge of matching scaffold degradation rates to the ability of cells to produce sufficient matrix, and also on the ability of the new tissue to respond appropriately to normal loading conditions upon implantation.86 4.2. Extension and Bending. Annular ligaments, as described previously, connect the cartilage rings of the trachea and enable longitudinal extension and bending to accommodate diaphragm contraction, lung inflation, and neck movements. Figure 6 shows a juvenile ovine trachea in resting

samples of different sizes. This normalization unfortunately eliminates the volume units, resulting in a compliance value with the less intuitive units cmH2O1−. Airtight experimental setups are required to measure compliance. Such setups have been used extensively in the study of native trachea.42,57,76,77 A review of the literature reveals a range of values for tracheal compliance, as shown in Figure 5.6,78−83 Typically, deflation curves are reported. Where publications reported inflation and deflation curves, only deflation curves were summarized. The publications by Croteau et al.78 and Bhutani et al.82 include additional age groups beyond what are plotted here. Some of these studies have evaluated preterm and full term lambs with and without trachealis muscle stimulation.6,79 Costantino et al. reported 0.013 and 0.053 cmH2O1− inflation and collapsing compliances (ΔP = 0 ± 10 cmH2O1−), respectively, in an in vitro study on preterm lambs.79 Transmural pressures have been shown to change progressively for 45 min after the trachea has been inflated or deflated, although it has been observed that 80% of the volume change occurs within the first 30 s.77 Costantino et al. reported that a 1 min equilibration time was sufficient for stable readings.79 The surrounding tissue in vivo plays a role in supporting the trachea, decreasing compliance as compared to ex vivo measurements.42 Tracheal compliance must be compared at given transmural pressures, as it is nonlinear. The trachea is more compliant under initial, small changes in transmural pressure, and then plateaus as transmural pressure increases. These characteristics make productive cough easy to achieve, while avoiding complete collapse. This is distinct from what would be seen in a uniform thick walled tubular structure, where compliance is initially low and then increases dramatically after collapse.84 While not physiological, understanding thick walled tubular collapse still provides some important insight into the design of constructs for trachea replacement. Importantly, the pressure gradient required for collapse of a thick walled tube is affected by the bending stiffness of the tube wall.84 Bending stiffness is modulated, in part, by the ratio of tube radius to wall thickness.84 This means that the actual dimensions of a developed tubular construct are a large factor in its mechanical ability to resist collapse. Even in a biomimetic U-shaped construct, after initial invagination of an elastomeric region, the latter portions of the compliance curve will still be defined by tube wall bending. Whether circumferentially uniform or Ushaped, the mechanical integrity of an original design will not be preserved if construct size is increased and no other adjustments are made. For this reason, working with animal models that have tracheae close in size to that of humans will be important for translating positive experimental results. Furthermore, it can be observed from Figure 5, that rabbits, a common small animal model for tracheal repair have significantly lower tracheal compliance at low transmural pressure. As compliance matching may not be a concern for trachea replacements in these animals, it brings the translational relevance of experimental results into question. For additional considerations with regards to animal model selection, refer to the review by ten Hallers et al.85 The ability of engineered or native tracheal tissue to deform radially, as can be assessed with cyclic lateral compression or compliance tests, has important implications for cartilage mechanical properties. A tissue engineered construct without the ability to deform in cross-section would not allow for regular mechanical loading of the cartilage. Cartilage is known to be mechanosensitive, and deteriorates in the absence of

Figure 6. Annular ligament provides extensibility to the trachea, which can be seen as the pink bands between the cartilage rings in an extended position. Images are of a single juvenile sheep specimen.

and extended positions. The native trachea has a very low bending stiffness, with no significant difference between the force required for forward vs left−right bending.87 Hoffman et al. measured the longitudinal composite elastic modulus of porcine trachea specimens containing bands of cartilage and annular ligament and reported a modulus of 1.10 ± 0.68 MPa, which was significantly different from the tensile elastic modulus for cartilage alone.51 We reported a tensile composite modulus for juvenile ovine trachea of 1.5 ± 0.37 MPa, and a toe region modulus of 0.10 ± 0.04 MPa.88 Behrend et al. evaluated the breaking strength of composite ovine trachea samples under longitudinal tension and found a mean breaking force of 198 N.89 They unfortunately did not present stress−strain data or a composite modulus. However, based on the information given regarding tracheal dimensions, we estimated a composite longitudinal toughness (also known as work of extension) between 85 and 240 kJ/m3. Toughness represents the energy dissipated in elastic and plastic deformation, as well as failure. Toughness, and particularly resilience (energy dissipated in elastic deformation), may be useful metrics for assessing engineered constructs. While numerous reviews on tracheal tissue engineering all call attention to the importance of longitudinal extensibility, the magnitude of normal extension in the native organ is often overlooked. During normal respiration, the healthy trachea may extend 20% in the adult and up to 46% in neonates.90 A G


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describing engineered tracheae continue to focus on materially uniform structures and lack extensive quantitative mechanical analysis. The following subsections discuss construct level mechanics of synthetic, decellularized, and tissue engineered trachea. 5.1. Synthetic Constructs. Solid tubular conduits were some of the first synthetic approaches attempted for trachea replacement. A large variety of materials was investigated for solid prosthesis ranging from extremely rigid, like stainless steel or glass,99,100 to flexible elastomer combinations.101 However, solid synthetic tubes are not incorporated well and typically dislodge, migrate, or become occluded.1 Solid prostheses additionally do not allow for the development of a blood supply to support epithelium. Attention then turned to the synthesis of porous constructs with the idea that they would better integrate with recipient tissues. In 2001, Kawaguchi et al. evaluated several models of acellular tubular mesh constructs with a spiral stent via radial compression at multiple loading rates and results normalized based on both the tube diameter and length.67 They compared these results to radial compression of dog trachea.67 Despite the presence of such testing methods in the literature, few publications in the subsequent 15 years performed mechanical analysis on tubular structures for trachea replacement, synthetic or tissue engineered. As stated in a 2002 review by Grillo, “the list of [material] combinations is almost endless.” While the list of investigated materials has continued to grow, none has yielded a consistently functional implant. What is clear from his review at that time is that porous prostheses (1) do show improved integration, in that dislodgement and migrations are no longer primary problems, (2) epithelialization in long segment acellular constructs remains incomplete, leading to infection, and (3) scar tissue formation and subsequent occlusion arise in most cases.1 As described in this review, we hypothesize that improved biomechanical properties in tracheal implants may help reduce this later concern of fibrotic stenosis. Scaffold mechanical characterization will be an important step toward the development of a successful solution. Ott et al. recently reported multiple mechanical analyses on several biodegradable polycaprolactone (PCL) and poly(lactic-coglycolic acid) (PLGA) electrospun scaffolds.102 They performed circumferential tensile, burst pressure, suture retention, and radial compression studies. They identified a gradient structure that yielded superior resistance to radial collapse; however, complete recovery was not observed upon load removal. Our group has also been working to develop polymeric scaffolds with physiologic mechanical properties. We evaluated tubular HEMA hydrogel constructs with varied ratios of hard to soft band widths under longitudinal tension, full radial compression, and cyclic partial radial compression at neonatal respiratory rates. Ultimately, we demonstrated that biomimetic patterning not only produced longitudinal extensibility but also simultaneously enhanced radial recoil.66 We further investigated construct compliance, and the role of tubular geometry in allowing for radial deformation of a physiological profile.88 5.2. Decellularized Trachea. Replacement of the trachea with decellularized tissue, in the absence of reseeding, has not been successful. In vivo evaluations of chemically processed tracheae have almost uniformly shown replacement of the decellularized structure with scar tissue and severe stenosis.103−105 Yet, as of 2016, decellularized cadaveric tissue was still reported to be the best clinically documented tracheal

longitudinally rigid tube will likely not allow for normal respiration.16 In a rare congenital anomaly, anterior tracheal cartilage can be vertically fused, leading to ongoing respiratory problems. In a case study describing four newborns, anterior vertical tracheal fusion was determined to be responsible for recurrent lower respiratory tract infections, reactive airway disease, and chronically retained secretions in all patients, likely due to modified airflow dynamics.91 These conditions were not secondary to surgical intervention or tracheal stenosis. Although the severity of these conditions decreased with age in the three surviving patients, observations of this anomaly indicate that changes in airway dynamics due to vertical fusion may predispose patients to chronic pulmonary disease. While patients with tracheal fusion are not necessarily candidates for trachea replacement, knowledge of the ramifications of these defects is pertinent to the development of replacements. With this in mind, development of a solid, cartilaginous, tissueengineered tube is not a suitable alternative. In the case of long segment defects, a prosthesis may provide additional length to reduce anastomotic tension 5 and accommodate the motions of respiration. However, inserting a rigid cartilage tube would cause significant longitudinal mechanical mismatch.16 Pulling at the attachment sites may lead to ongoing inflammation and fibrosis. As discussed in Section 3, increased airway resistance due to fibrotic stenosis could lead to abnormal loading conditions, resulting in collapse of an otherwise mechanically sufficient implant. In addition to potentially prolonging inflammation due to mechanical mismatch, a rigid tube will not allow for normal stretching of the tracheal epithelium. Tracheal extension likely plays a role in regulation of the airway epithelium and its immune and secretory functions. Numerous studies in multiple species have demonstrated that tracheal epithelium is mechanosensitive, responding to tensile and compressive forces.92−95 Physiological levels of cyclic strain have been shown to down-regulate eicosanoid synthesis in a frequency dependent manner in both cat tracheal epithelium and Calu-3 cells (derived from human lung adenocarcinoma, but shown to resemble tracheal epithelial cells).92 Eicosanoids help regulate smooth muscle tone, vascular permeability, mucous secretion, and immune function.96,97 Tschumperlin and Drazen published a review on the role of mechanical stimuli in airway remodeling98 and concluded that epithelial cells produce factors that modulate the inflammatory environment of the airway wall, as well as fibroblast recruitment and proliferation. It is notable that either the absence or excess of mechanical loading on tracheal epithelium may have negative consequences with regards to inflammation mediated fibrosis.

5. TRACHEA REPLACEMENT STRATEGIES AND MECHANICS Tracheal collapse is recognized as a main mode of failure for trachea implants, with nearly all investigated materials showing inadequate circumferential strength.14 However, even in cases where mechanical strength may be sufficient, stenosis is a serious problem.7,8 As discussed in the previous sections, inadequate extensibility and compliance may exacerbate stenotic fibrosis, promote mucous accumulation and thereby infection, limit mechanoregulation of engineered tissues, and produce abnormal loading conditions; combined, these outcomes all contribute to construct collapse. Despite repeated urging in many reviews that longitudinal extensibility and radial rigidity will be required in a trachea replacement, publications H


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ACS Biomaterials Science & Engineering implant.106 Cadaveric allografts have almost all ultimately required long-term stenting due to structural instability,107−109 possibly due to the absence of live chondrocytes maintaining the matrix.107 Another important limitation of fixed cadaver tissue is that it would not be expected to grow along with pediatric patients. Recellularization of decellularized constructs, i.e. the use of decellularized tissue as scaffolds for tissueengineered replacements, is promising. Part of the rationale behind using decellularized scaffolds is that they maintain their native mechanical properties.110 However, Haykal et al. investigated three different trachea decellularization protocols and found increased compliance, or risk for collapse, with all methods.81 Partington et al. subsequently suggested that more work is needed to identify decellularization protocols that maintain tissue tensile strength, as they observed significant weakening with decellularization that could pose risk for collapse.111 Investigation of longitudinal tensile properties of decellularized human trachea found a toeregion modulus of 0.23 ± 0.03 MPa, and a high strain tensile modulus of 4.4 ± 2.0 MPa,112 which is higher than values reported for fresh porcine trachea.51 However, a loss of mechanical integrity was observed over time in stored decellularized trachea that could not be ameliorated with genipin cross-linking. On the basis of these findings, mechanical analysis should be a key aspect of publications on decellularized and recellularized grafts. 5.3. Tissue Engineered. Focus has turned to tissue engineering for a trachea replacement solution. Several reviews9−13 have chronicled the trends in pediatric and adult tracheal replacement over the past 10 years (2006−2016) and have collectively emphasized the need for epithelialization, vascularization (particularly to support the epithelium), and addressing stenosis. Particularly for pediatric applications, tissue engineered solutions are preferred over nondegradable synthetic solutions because of the potential for growth. Reports of tissue-engineered replacements often quantify sulfated glycosaminoglycan content, as this is correlated with the mechanical integrity of native cartilage, and suggest that constructs are mechanically sufficient. Others claim that constructs have mechanical properties similar to native trachea, but present a qualitative assessment of bending and compressing with forceps or have not reported normalized results.113,114 Notably, Kojima et al. developed trachea constructs from nasal chondrocytes and found that the tensile modulus of the engineered cartilage was only 15% of native, despite similar GAG and collagen content. Biochemical assessment (e.g., GAG content, collagen content, etc.) cannot take the place of direct evaluation of mechanical properties. Indeed, most studies that have investigated material properties have found that they were significantly inferior to native tracheal cartilage.53,115−118 Beyond assessing material properties of engineered tissues, it is important to consider overall mechanical behavior at the construct level. To our knowledge, only two studies of biosynthetic tracheae from outside our group have reported quantitative analysis of longitudinal mechanical properties in any capacity.114,119 Both of these studies used three-point bending, which would reflect flexibility, but not directly construct extensibility. The finite element analysis of construct bending reported by Park et al. also did not reflect the mechanical contributions of the cellularization of their construct, which significantly modified the construct geometry. Our recent work reports composite moduli for constructs

evaluated under longitudinal tensile loading, and compares this to the composite moduli of ovine tracheae.88 This method is straightforward, both experimentally and computationally. We recommend direct longitudinal evaluation of scaffolds and tissue engineered prostheses for tracheal repair. Reporting of representative stress−strain curves would further enable crossstudy comparisons.

6. SUMMARY OF RECOMMENDED TESTS AND KEY CONSIDERATIONS To sufficiently evaluate engineered tracheae, and to better interpret the, often variable, results from in vivo studies, we recommend routine longitudinal tensile, lateral (radial) compression, and compliance testing. We here outline key considerations for each of the aforementioned testing strategies: Longitudinal testing provides information about tracheal extensibility, which can be compared to native values. When reporting longitudinal tensile data, loads should be normalized by specimen cross-sectional area. For full tracheae, this is the circumference multiplied by the wall thickness. We recommend using the term “normalized load” as opposed to stress. While informative, this value does not represent the true stress in any portion of a composite sample. Longitudinal displacements should be normalized by sample length, and termed “composite strain”. While informative, this value does not represent the stress tensor in any portion of a composite sample. We further suggest that moduli should be referred to as “composite moduli” for the aforementioned reasons. Composite moduli should be compared at low (toe region) and high strains, not only over the linear region of the normalized load vs composite strain curves for native tissues. The composite yield strain should be reported for engineered constructs, and resilience may also be informative. Lateral (radial) tubular compression, while not a typical loading regime in vivo, is significantly more straightforward to perform than compliance testing, and offers quantitative information on tubular resistance to collapse. Displacement values from lateral compression tests should be normalized by tube diameter, and force values should be normalized by tube length. Cyclic tests may also be useful for evaluating tube recoil. These tests should be performed at normal and elevated respiratory rates for relevance. Compliance tests are able to directly evaluate the ability of a trachea construct to withstand and vary with physiologic transmural pressures. When performing compliance tests, the sample length should be carefully considered to avoid edge effects. Short constructs may demonstrate artificially low compliance. Additionally, the change in volume at a given transmural pressure can be normalized by original internal volume to facilitate comparisons across samples and studies. 7. CONCLUSIONS The mechanical anisotropy of the trachea is essential to its function. Longitudinal extension enables normal respiration. Radial deformation enables productive coughing, while radial rigidity and recoil maintain a patent airway. Together, longitudinal and radial deformation have downstream implications for mechanotransduction in tracheal tissues, which plays a role in regulating inflammatory and immune responses, as well as maintaining tissue mechanics. I


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(14) Tan, Q.; Steiner, R.; Hoerstrup, S. P.; Weder, W. Tissueengineered trachea: History, problems and the future. European journal of cardio-thoracic surgery: official journal of the European Association for Cardio-thoracic Surgery 2006, 30 (5), 782−786. (15) Belsey, R. Resection and reconstruction of the intrathoracic trachea. Br. J. Surg. 1950, 38 (150), 200−5. (16) Macklin, C. C. The Musculature of the Bronchi and Lungs: A Retrospect. Can. Med. Assoc J. 1929, 20 (4), 404. (17) Bouhuys, A. The Physiology of Breathing: A Textbook for Medical Students; Grune & Stratton: New York, 1977; p 352. (18) Wailoo, M. P.; Emery, J. L. Normal growth and development of the trachea. Thorax 1982, 37 (8), 584−7. (19) Donaldson, S. W.; Tompsett, A. C., Jr Tracheal diameter in the normal newborn infant. American journal of roentgenology, radium therapy, and nuclear medicine 1952, 67 (5), 785−787. (20) Butz, R. O., Jr. Length and cross-section growth patterns in the human trachea. Pediatrics 1968, 42 (2), 336−341. (21) Teng, Z.; Trabelsi, O.; Ochoa, I.; He, J.; Gillard, J. H.; Doblare, M. Anisotropic material behaviours of soft tissues in human trachea: an experimental study. J. Biomech. 2012, 45 (9), 1717−1723. (22) Safshekan, F.; Tafazzoli-Shadpour, M.; Abdouss, M.; Shadmehr, M. Mechanical Characterization and Constitutive Modeling of Human Trachea: Age and Gender Dependency. Materials 2016, 9 (12), 456. (23) Wachsmuth, L.; Soder, S.; Fan, Z.; Finger, F.; Aigner, T. Immunolocalization of matrix proteins in different human cartilage subtypes. Histol Histopathol 2006, 21 (5), 477−485. (24) Kelly, D. J.; Crawford, A.; Dickinson, S. C.; Sims, T. J.; Mundy, J.; Hollander, A. P.; Prendergast, P. J.; Hatton, P. V. Biochemical markers of the mechanical quality of engineered hyaline cartilage. J. Mater. Sci.: Mater. Med. 2007, 18 (2), 273−81. (25) Knudson, C. B.; Knudson, W. Cartilage proteoglycans. Semin. Cell Dev. Biol. 2001, 12 (2), 69−78. (26) Hinderer, S.; Schesny, M.; Bayrak, A.; Ibold, B.; Hampel, M.; Walles, T.; Stock, U. A.; Seifert, M.; Schenke-Layland, K. Engineering of fibrillar decorin matrices for a tissue-engineered trachea. Biomaterials 2012, 33 (21), 5259−66. (27) Bianco, P.; Fisher, L. W.; Young, M. F.; Termine, J. D.; Robey, P. G. Expression and localization of the two small proteoglycans biglycan and decorin in developing human skeletal and non-skeletal tissues. J. Histochem. Cytochem. 1990, 38 (11), 1549−63. (28) Plaas, A. H.; Neame, P. J.; Nivens, C. M.; Reiss, L. Identification of the keratan sulfate attachment sites on bovine fibromodulin. J. Biol. Chem. 1990, 265 (33), 20634−20640. (29) Melching, L. I.; Roughley, P. J. Modulation of keratan sulfate synthesis on lumican by the action of cytokines on human articular chondrocytes. Matrix Biol. 1999, 18 (4), 381−90. (30) Johnson, H. J.; Rosenberg, L.; Choi, H. U.; Garza, S.; Hook, M.; Neame, P. J. Characterization of epiphycan, a small proteoglycan with a leucine-rich repeat core protein. J. Biol. Chem. 1997, 272 (30), 18709−17. (31) SundarRaj, N.; Fite, D.; Ledbetter, S.; Chakravarti, S.; Hassell, J. R. Perlecan is a component of cartilage matrix and promotes chondrocyte attachment. J. Cell Sci. 1995, 108 (Pt 7), 2663−2672. (32) Mow, V. C.; Kuei, S. C.; Lai, W. M.; Armstrong, C. G. Biphasic creep and stress relaxation of articular cartilage in compression? Theory and experiments. J. Biomech. Eng. 1980, 102 (1), 73−84. (33) Maroudas, A. Distribution and diffusion of solutes in articular cartilage. Biophys. J. 1970, 10 (5), 365−379. (34) Schulz, R. M.; Bader, A. Cartilage tissue engineering and bioreactor systems for the cultivation and stimulation of chondrocytes. Eur. Biophys. J. 2007, 36 (4−5), 539−68. (35) Rains, J. K.; Bert, J. L.; Roberts, C. R.; Pare, P. D. Mechanical properties of human tracheal cartilage. J. Appl. Physiol. 1992, 72 (1), 219−25. (36) Guyton, A. C.; Hall, J. E. Textbook of Medical Physiology, 11th ed.; Elsevier Saunders: Philadelphia, 2006; p 1116. (37) Teng, Z.; Ochoa, I.; Li, Z.; Lin, Y.; Rodriguez, J. F.; Bea, J. A.; Doblare, M. Nonlinear mechanical property of tracheal cartilage: a

Despite the clear importance of mechanics in tracheal function, efforts to produce a trachea replacement have continued to focus primarily on the evaluation of additional scaffold materials and culture conditions. There is a need for clear design criteria, and straightforward, quantitative mechanical metrics at the construct level that are regularly reported across studies. Direct mechanical evaluation of tubular conduits, as opposed to determination of material properties, may be useful for demonstrating mechanical sufficiency of engineered constructs, as the tube radius to thickness ratio affects compliance and collapse. Furthermore, it is unlikely that one tissue type will enable longitudinal extension and provide sufficient radial rigidity. Increased attention to the complexity of the trachea as an organ with multiple tissue types, as opposed to an engineered cartilaginous conduit, may lead to progress in the development of a successful tissue engineered trachea.

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AUTHOR INFORMATION

Corresponding Author

*E-mail: [email protected]; Phone: 617-373-6243. ORCID

Debra T. Auguste: 0000-0002-7389-9331 Notes

The authors declare no competing financial interest.

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REFERENCES

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