Effect of polymer fabrication (UV exposure) and postprocessing (sanding) on uniform QHM polymer (nongraded) porosity, roughness, and hydrophobicity (contact angle)
Prior to in vitro studies, it is crucial to determine the effect of polymer fabrication and postprocessing steps on QHM polymer properties, such as porosity, pore size, surface roughness, and hydrophobicity (contact angle), as these biomaterial attributes can regulate cell behavior and matrix remodeling23,24,25. Polymer fabrication procedures such as UV light-induced crosslinking showed minor differences in QHM polymer porosity, with 0 s UV and 300 s UV QHM polymers exhibiting 10 and 7% porosity, respectively (Fig. 2A, B). In general, increased UV exposure time resulted in reduced porosity, but pore sizes remained similar at ~0.008 µm (Fig. 2B). Polymer postprocessing procedures such as sanding resulted in a similar surface roughness of ~4 µm based on computed arithmetic calculations (Fig. 2C, D). Additionally, QHM polymers exhibited contact angles of less than 90° for all samples, demonstrating hydrophilicity and suitability for subsequent GF biopatterning, with only the 0 s UV QHM polymer exhibiting a slightly reduced contact angle (by ~10°) relative to those of the 90 s UV, 180 s UV, and 300 s UV QHM polymers (Fig. 2E, F). Thus, QHM polymer fabrication and postprocessing did not change pore size or surface roughness but had a minor impact on porosity and hydrophilicity.
A Schematic diagram depicting mercury intrusion porosimetry. B Average porosity and pore size of QHM polymers. Porosity decreased by less than 2.5% with increasing UV exposure (n = 3). Average pore diameter of QHM polymers. Pore size was not altered by UV exposure (n = 3). C Schematic diagram depicting line profilometry. D Surface roughness of QHM polymers. The computed arithmetic average roughness remained unchanged by sanding (n = 3). E Schematic diagram depicting hydrophobicity (contact angle) measurements. F Hydrophobicity (contact angle) of QHM polymers. The contact angle increased by 10° with increased UV exposure between the 0 s UV with 90 s UV QHM polymer and remained unaltered thereafter (n = 6). Error bars indicate the standard error of the mean. Statistical significance as indicated.
Effect of polymer fabrication (UV exposure) on mechanically graded QHM polymer material stiffness (microscale elastic modulus)
To determine if the QHM polymer can mimic the bone-tendon interface within physiologically relevant values, i.e., length of the human bone-tendon interface: 0.1–1 mm10, mechanically graded QHM polymers with 4 different graded regions spanning 1–2 mm were fabricated. In general, a trend of increased microscale stiffness with increased UV exposure time was observed (Fig. 3 and Fig. S1). Thus, QHM polymers with controlled UV exposure can be fabricated as mechanically graded materials that mimic the physiological dimensions of the bone-tendon interface.
A Schematic diagram depicting the fabrication of mechanically graded QHM with graded regions of 250 and 500 μm. B Normalized elastic modulus of mechanically graded QHM. Graded regions spanning 250 μm (n = 12) and 500 μm (n = 6) in width were fabricated, and an increasing gradient of elastic moduli was observed with increased UV exposure. Error bars indicate the standard error of the mean; statistical significance. *p ≤ 0.05.
Effect of material stiffness and GFs (nonpatterned soluble and biopatterned) on osteoblast and tenocyte differentiation in vitro
To determine the combinatorial effect of QHM material stiffness (as a macroscale biomechanical cue) and GFs (as biochemical cues) on bone and tendon cell differentiation, C2C12 cells, which are a myoblast cell line capable of differentiating into osteoblasts and tenocytes16,17, were cultured in the presence of bone- and tendon-promoting GFs on various QHM polymers. Different UV exposure times (0, 90, 180, and 300 s) resulted in QHM polymers with macroscale tensile modulus values of 0.6, 1.7, 2.5, and 2.7 GPa, respectively, as described in our prior work9 and here (Fig. S2). Such mechanical properties approximate the tensile modulus of the human supraspinatus tendon (0.2–0.6 GPa)1 and cortical bone (11.0–29.0 GPa)14.
Initially, soluble (i.e., nonpatterned) GFs were added to the culture medium to ensure that the cells in different QHM polymer groups received similar dosages, thus eliminating any potential confounding variables as a result of differential GF binding to patterned substrates. ALP staining and von Kossa staining were used to test osteogenic differentiation, and IF staining (for the markers SCX, TNMD, and TNC) was used to test tenogenic differentiation. BMP-2 was used as a bone-promoting GF cue based on previous studies15,16,17. When cultured in the presence of BMP-2 for 4 and 6 days, C2C12 cells exhibited increased activity of the osteoblast marker ALP on stiffer QHM polymers (Fig. 4A). With further culture (27 days), comparable C2C12 mineralization, as indicated by von Kossa staining, was observed on QHM polymers in all groups (Fig. 4B). It is worth noting that BMP-2 can also induce chondrogenesis with increased ALP activity21, which is not discussed here. Thus, QHM biomaterial macroscale stiffness has an early effect on BMP-2-mediated C2C12 osteoblast-like differentiation.
C2C12 cells were cultured on QHM polymers with different material stiffnesses (as a result of different UV exposure) at designated time points. A After 4–6 days of culture in either soluble 100 ng/mL or 200 ng/mL BMP-2-containing medium, increased ALP staining of C2C12 cells was observed on QHM polymers with increased substrate stiffness. This trend was impacted by the duration (4 days BMP-2 + 2 days no BMP-2 versus 6 days BMP-2) of BMP-2 administration (n = 3). B After 27 days of culture, comparable C2C12 cell-mediated mineralization indicated by von Kossa staining was observed on QHM polymers in all groups (n = 6–9). ALP-positive regions are stained blue. von Kossa-positive regions are stained black (black arrows). Scale bars are as indicated. TCPS, tissue culture-grade polystyrene. Error bars indicate the standard error of the mean; statistical significance. *p ≤ 0.05.
Material stiffness also affected C2C12 tenocyte differentiation. FGF-2 was used as a tendon-promoting GF cue based on previous studies16,17. In our study, FGF-2 increased the expression of TNC and TNMD of C2C12 cells in a dose-dependent manner when the cells were cultured on TCPS for 3 days, and 100 ng/mL FGF-2 showed increased C2C12 tenogenic differentiation relative to that of the 0 ng/mL and 50 ng/mL FGF-2 groups. When cultured in the presence of FGF-2 for 3 days, C2C12 cells exhibited decreased expression of the tendon marker Scleraxis (SCX) on stiffer QHM polymers (Fig. 5). Thus, QHM biomaterial macroscale stiffness has an early effect on FGF-2-mediated C2C12 tenocyte-like differentiation.
A FGF-2 increased the expression of TNC and TNMD in C2C12 cells in a dose-dependent manner when the cells were cultured on tissue culture-grade polystyrene. B C2C12 cells were cultured on QHM polymers with different material stiffnesses (as a result of different UV exposure). After 3 days of culture in 100 ng/mL FGF-2-containing medium, decreased SCX expression was observed with increased substrate stiffness (n = 9). TNC-, TNMD-, or SCX-positive regions are shown in white; scale bars are as indicated; error bars indicate the standard error of the mean; statistical significance. *p ≤ 0.05.
Our data (Figs. 4 and 5) showed that soluble or liquid-phase GFs and biomechanical cues interact. However, some tissue engineering strategies utilize GF-containing scaffolds, where biochemical cues may be conjugated or immobilized in a solid-phase form to facilitate persistent signaling. As such, we aimed to study whether a similar phenomenon occurred with solid-phase biochemical cues interacting with biomechanical cues. Prior to such studies, we first determined GF retention on the scaffolds after inkjet biopatterning15,16,17,18,20,21. QHM polymers patterned with either BMP-2 or FGF-2 retained pattern fidelity after multiple PBS washes. Although GF desorption was observed following the first PBS wash, no further losses occurred with additional washes (Fig. 6A, B). This demonstrated the persistence of biopatterned GFs, which is crucial for long-term spatial control of musculoskeletal differentiation and has been demonstrated in our prior in vivo studies15,21. Furthermore, shelf-life studies demonstrated that GF bioactivity was retained under extended storage conditions at 4 °C, with the majority of GF-biopatterned QHM polymers retaining their ability to spatially direct osteoblast differentiation 3 months post-printing (Fig. S3).
A BMP-2 immobilization on fibrin-coated QHM polymers. B FGF-2 immobilization on fibrin-coated QHM polymers. Successive washing did not alter the amount of fluorescently labeled BMP-2 or FGF-2, demonstrating persistent immobilization (n = 3). Fluorescence intensity was normalized to sample images acquired prior to washing; error bars indicate the standard error of the mean.
The effect of GF biopatterning and QHM polymer material stiffness on bone and tendon cell differentiation was also examined. BMP-2 and FGF-2 biopatterning on fibrin-coated QHM increased C2C12 ALP staining and SCX expression relative to that of their respective off-pattern controls. Additionally, similar to soluble GF studies, GF patterns increased ALP activity and decreased SCX expression in C2C12 cells with increased material stiffness. The interaction effects were significant for different material stiffnesses and biopatterned BMP-2 on ALP activity (F(3,40) = 3.127, p = 0.036, partial η2 = 0.190), as well as for different material stiffness and biopatterned FGF-2 on SCX expression (F(3,40) = 6.428, p = 0.001, partial η2 = 0.325; Fig. 7). A similar trend was also observed in C3H10T1/2 cells—increased ALP staining was found on stiffer materials within BMP-2 biopatterned regions (Fig. S4).
A Schematic diagram depicting the experimental setup of in vitro BMP-2 or FGF-2 biopatterning studies. B Effect of BMP-2 biopatterning and QHM material stiffness on the spatial control of C2C12 osteoblast differentiation. Increased ALP staining was observed on BMP-2 patterned regions relative to the nonpatterned control as well as on BMP-2 patterned regions of QHM polymers that exhibit increased substrate stiffness (relative to the 0 s UV QHM polymer; n = 6). C Effect of FGF-2 biopatterning and QHM material stiffness on the spatial control of C2C12 tenocyte differentiation. Increased SCX expression was observed in FGF-2 patterned regions relative to the nonpatterned control. Decreased SCX expression was observed on FGF-2 patterned regions of QHM polymer that exhibit increased substrate stiffness (relative to the 0 s UV QHM polymer; n = 6). ALP-positive regions are stained blue, and SCX-positive regions are shown in white. Scale bars are as indicated. Error bars indicate the standard error of the mean; statistical significance. *p ≤ 0.05.
Altogether, these data showed that increased QHM polymer stiffness as a result of longer UV exposure (i.e., macroscale tensile moduli: 0.6 GPa–2.7 GPa based on 0–300 s UV exposure time)9, increased BMP-2-mediated C2C12 osteoblast differentiation and decreased FGF-2-mediated tenocyte differentiation, regardless of whether the GF cue was soluble or biopatterned. Thus, GF(s) and material stiffness can be used together to spatially control musculoskeletal cell differentiation in vitro.
Effect of blebbistatin (a mechanosensing inhibitor) on material stiffness- and GF-mediated osteoblast differentiation
To further determine the involvement of the material stiffness of the QHM polymer on C2C12 osteoblast differentiation, blebbistatin26, which disrupts mechanosensing, was utilized together with soluble GFs in subsequent experiments. As such, C2C12 cells were cultured on QHM polymers with various tensile moduli in the presence of BMP-2, with or without 25 μM blebbistatin (Fig. 8A).
A Schematic diagram depicting the experimental setup of in vitro mechanosensing studies. B Effect of QHM polymer stiffness on soluble BMP-2-mediated C2C12 osteoblast differentiation after 4, 8, and 14 days of culture in the presence or absence of 25 μM blebbistatin. After 4 days of culture, increased ALP staining was observed with increased substrate stiffness. Blebbistatin inhibited substrate stiffness-mediated ALP staining, but the overall trend of increased ALP staining with increasing substrate stiffness remained. Blebbistatin-mediated inhibitory effects were more obvious on less stiff QHM polymers and less obvious after longer culture periods. The ALP inhibitory index was defined as the ratio of ALP activity for blebbistatin-treated cells to DMSO-treated cells; a value of 1.0 indicated no inhibition, while a value of 0.0 indicated complete inhibition (n = 3). ALP-positive regions are stained blue. TCPS, tissue culture-grade polystyrene. Error bars indicate the standard error of the mean; statistical significance. *p ≤ 0.05.
Our data showed that BMP-2-induced ALP activity was slightly inhibited by blebbistatin supplementation, although the overall trend of increased ALP activity on stiffer QHM polymers remained (F(1,20) = 98.559, p ≤ 0.001, partial η2 = 0.831). Indeed, the interaction effect of different material stiffnesses and blebbistatin treatment on BMP-2-induced ALP activity was significant (F(4,20) = 14.768, p ≤ 0.001, partial η2 = 0.747). The level of blebbistatin-induced ALP inhibition was determined by calculating the ratio of ALP activity between the 25 µM blebbistatin-treated group and its corresponding 0.3% DMSO vehicle control (ALP inhibitory index values of 0.0 and 1.0 indicate complete inhibition and no inhibition, respectively). Analysis indicated that a higher level of C2C12 ALP inhibition was observed on less stiff 0 s UV QHM polymers, e.g., C2C12 ALP inhibition in 0 s UV QHM was higher than that in 300 s UV QHM, demonstrating that cells on less stiff materials were more sensitive to mechanosensing than cells on stiffer materials. With increased culture duration, inhibition of ALP activity was observed only on the least stiff QHM polymer (0 s UV) at 8 days and abrogated by 14 days (Fig. 8B). These results suggested that material stiffness had an obvious effect on the early but not the late phase of BMP-2-mediated osteoblast differentiation. The eventual osteogenic differentiation and mineralization of cells on all QHM polymers may be explained by the prolonged stimulatory effects of BMP-2, which eventually overcomes material stiffness effects, as an increased frequency of BMP-2 administration or an increased dose also increased ALP activity (Fig. 4A). Thus, the varying levels of osteoblast and tenocyte differentiation observed during culture on QHM polymers of varying stiffness in vitro were attributed to mechanosensing effects.
Effect of GF-biopatterned, mechanically graded (biphasic) QHM polymers on musculoskeletal differentiation in vitro and in an ectopic, mouse subcutaneous implantation model
Having established that material stiffness and GFs interact to affect musculoskeletal differentiation on uniform, nongraded QHM polymers, we next assessed the combinatorial effect of mechanically graded QHM polymers and GFs on bone and tendon cell differentiation in vivo. To simplify the study design, a biphasic bone-tendon scaffold consisting of only 0 s UV and 300 s UV QHM polymer regions was used. Owing to its phototunable properties, this biphasic material approximates the tensile attributes of the human supraspinatus tendon and cortical bone9, which is most commonly torn in rotator cuff tears1.
Similar to prior in vitro culture conditions, C2C12 cells were cultured on mechanically graded QHM polymers in the presence of liquid-phase bone- or tendon-promoting GFs, i.e., BMP-2 and FGF-2, respectively. Our data showed that the stiffer (300 s UV) region of QHM polymers showed increased ALP staining but decreased SCX expression relative to less stiff (0 s UV) regions and vice versa (Fig. S5), which was similar to the results for nongraded QHM polymers.
Subcutaneous implantation of biphasic, fibrin-coated QHM polymers patterned with bone- and tendon-promoting cues in mice demonstrated biocompatibility, as well as the formation of bone- and tendon-like tissues relative to nonprinted controls. After implantation, the mice showed consistent weight gain and did not exhibit adverse clinical signs or mortality, including detrimental reactions such as necrosis, infection, and granuloma around the implants (data not shown). Scaffold regions biopatterned with BMP-2 (300 s UV QHM) resulted in bone-like tissue with the formation of collagen-rich structures (shown in trichrome and hematoxylin & eosin/H&E staining) that contained abundant bone marrow and bone cells, including multinucleated, tartrate-resistant acid phosphatase/TRAP-positive osteoclasts, as well as cuboid-shaped osteoblasts that lined the bone surface. Scaffold regions biopatterned with FGF-2 (0 s UV QHM polymer) resulted in tendon-like tissues with birefringent (shown in polarized light microscopy), wavy and crimped collagen fibers (shown in trichrome and H&E staining) that contained cells that strongly expressed the tenocyte marker SCX (Fig. 9 and S6). Altogether, these data demonstrated that bone- and tendon-promoting GF biopatterning along with biphasic scaffolds with bone- and tendon-like macroscale biomechanical properties spatially controlled ectopic bone- and tendon-like tissue formation in vivo.
A Schematic diagram depicting the experimental setup of in vivo BMP-2 and FGF-2 biopatterning studies. B Effect of BMP-2 and FGF-2 biopatterning on the spatial control of bone- and tendon-like differentiation after 14 days of subcutaneous implantation in mice. SCX-positive regions containing birefringent and wavy collagen fibers reminiscent of tendon-like tissue were observed on QHM polymers patterned with FGF-2; TRAP-positive, collagen-rich extracellular structures that contained bone cells and bone-marrow tissue reminiscent of bone-like tissue were observed on QHM polymers patterned with BMP-2 (n = 9). Scale bars are as indicated.