Crosslinking kinetics of single-network PAM hydrogelsTo be suitable for DASP printing, the bio-ink should possess a gelation time of sufficient duration. This allows the printed droplets to undergo partial swelling, enabling them to partially coalesce with adjacent droplets before complete crosslinking occurs (Fig. 1a). However, the crosslinking time should not be too long, otherwise the bio-ink may diffuse through the porous supporting matrix, leading to uncontrollable droplet shape8. Thus, it is essential to quantify the crosslinking kinetics of single-network PAM hydrogels.Unlike typical hydrogels formed by crosslinking precursor polymer chains using small molecules as crosslinking agents, our single-network PAM gels are formed by crosslinking a pair of linear PAM polymers. The two polymers are of the same molecular weight (MW) but are functionalized by tetrazine (TZ) (Supplementary Fig. 3) and norbornene (NB) groups, respectively. Upon reaction, the TZ and NB groups form a covalent bond via additive-free click reaction at room temperature (RT) (Fig. 1b, i). Therefore, the crosslinking kinetics of our single-network PAM hydrogels is expected to depend on both the concentration, \(c\), of the polymers and the grafting ratio, \({f}_{x}\), of the functional groups, which is defined as the molar ratio between the grafted functional groups and the total number of chemical monomers on the PAM backbone.To investigate the crosslinking kinetics of PAM hydrogels, we maintain the concentrations of TZ-PAM and NB-PAM solutions at 10% (w/v) before mixing while varying the molar ratio between NB and TZ groups. To do so, we keep the grafting ratio of TZ-PAM, \({f}_{x,{TZ}}\), at 2.53% (Supplementary Fig. 4); on average, this value corresponds to approximately 50 TZ groups per polymer. We vary the grafting ratio of NB groups, \({f}_{x,{NB}}\), from 0.63% to 4.25% (Supplementary Fig. 5). These formulations cover NB/TZ ratios spanning approximately one order of magnitude, ranging from 0.25 to 1.68. For each mixture with equal amount of TZ-PAM and NB-PAM polymers, we employ a stress-controlled rheometer to continuously monitor its viscoelasticity in real time for 12 h (see Methods). We determine the gelation time, \({t}_{g}\), above which the storage modulus, \(G^{{\prime}}\), exceeds the loss modulus, \(G^{{\prime} {\prime}}\), as denoted by the arrow in Fig. 2a. As the NB/TZ ratio increases from 0.25 to 1.68, \({t}_{g}\) decreases by nearly 40 times from 616 s to 16 s (solid squares in Fig. 2b).Fig. 2: Crosslinking kinetics and stiffness of single-network PAM hydrogels crosslinked through biofriendly click chemistry.a Real-time characterization of storage (\(G^{\prime}\), thick lines) and loss (\(G^{{\prime} {\prime}}\), thin lines) moduli for mixtures of NB-PAM and TZ-PAM at various molar ratios between NB and TZ groups. The concentrations of TZ-PAM and NB-PAM solutions are both fixed at 10% (w/v) before mixing. For TZ-PAM, the grafting ratio of TZ is fixed at 2.53%, defined as the molar ratio between the grafted tetrazine and the total number of chemical monomers on the PAM backbone. This value corresponds to on average of 52 TZ groups per polymer. In contrast, for NB-PAM, the grafting ratio of NB groups varies from 0.63 to 4.25%. The gelation time, \({t}_{g}\), is defined as the point above which \(G^{\prime}\) surpasses \(G^{{\prime} {\prime}}\). All measurements are conducted at RT, using a fixed strain 0.5% and an oscillatory shear frequency 1 Hz. b The dependencies of gelation time, \({t}_{g}\), on the NB/TZ ratio at fixed polymer concentration of 10% (w/v) and on the polymer concentration, \(c\), at fixed NB/TZ ratio of 1. c The dependencies of storage modulus, \(G^{{\prime}}\), at gelation time, \({t}_{g}\), on the NB/TZ ratio at fixed polymer concentration of 10% (w/v) and on the polymer concentration, \(c\), at fixed NB/TZ ratio of 1. d The dependencies of storage (\(G^{\prime}\), thick lines) and loss (\(G^{{\prime} {\prime}}\), thin lines) moduli of a completely crosslinked single-network PAM hydrogels on the oscillatory shear frequency. In all hydrogels, both the TZ-PAM and NB-PAM polymers have roughly the same grafting ratio of 2.5%, and they are mixed at the same concentrations ranging from 1.25 to 15% (w/v). e The dependencies of the experimentally measured (solid squares) and the theoretically predicted (dash line) shear modulus \(G\) for PAM hydrogels on the concentration, \(c\), of PAM polymers. \({c}^{*}\), overlap concentration of PAM solutions; \({c}^{*\ast }\), the crossover concentration at which the correlation length, \(\xi\) is about the average distance between two neighboring functional groups on the same the polymer chain, \({d}_{x}\). For \(c \, > \,{c}^{*\ast }\), \(\xi\, < \,{d}_{x}\). f Schematic of our theory explaining the relation between the concentration of PAM polymers and the network shear modulus. In a crosslinked network, the mesh size \({a}_{x}\) exhibits two distinct regimes: (1) for \(c \, < \,{c}^{*\ast }\), \({a}_{x}\,\approx \,\xi\); (2) for \(c \, > \,{c}^{*\ast }\), \({a}_{x}\) becomes saturated and \({a}_{x}\approx {d}_{x}\). The network modulus \(G\,\approx \,{k}_{B}T/{a}_{x}^{3}\), as indicated by the dashed line in (e).The decrease in the gelation time with the increase of NB/TZ ratio can be explained by a simple kinetic theory. The rate of forming crosslinks, \({{{{{\rm{d}}}}}}{C}_{x}/{{{{{{\rm{d}}}}}}t}\), is proportional to the probability of a NB group to meet with a TZ group, which is the product of the concentrations, \(({C}_{{NB},0}-{C}_{x})\) and \(({C}_{{TZ},0}-{C}_{x})\), of the two in the mixture:$$\frac{{{d}}{C}_{x}}{{{{d}}t}}=k\left({C}_{{NB},0}-{C}_{x}\right)({C}_{{TZ},0}-{C}_{x})$$
(1)
Here, k is a constant determined by the reaction rate between a NB and a TZ, and \({C}_{{NB},0}\) and \({C}_{{TZ},0}\), respectively, are the concentrations of unreacted NB and TZ groups at reaction time \(t=0\). Integrating Eq. (1), one obtains the relation between the concentration of crosslinks and reaction time:$$t=\frac{1}{k}\left\{\begin{array}{cc}\frac{1}{\left({C}_{{NB},0}-{C}_{{TZ},0}\right)}{ln}\left[\frac{({C}_{{NB},0}-{C}_{x}){C}_{{TZ},0}}{({C}_{{TZ},0}-{C}_{x}){C}_{{NB},0}}\right],&{C}_{{NB},0}\,\ne\, {C}_{{TZ},0}\hfill\\ \frac{{C}_{x}}{{C}_{0}\left({C}_{0}-{C}_{x}\right)},\hfill&{C}_{{NB},0}\,=\,{C}_{{TZ},0}\,=\,{C}_{0}\end{array}\right.$$
(2)
At the gelation time \({t}_{g}\), each polymer has on average one crosslink, forming a giant molecule that percolates the whole volume of the solution15. As each polymer carries about 50 functional groups, the concentration of crosslinks relative to the initial concentration of the functional groups, \(\beta \equiv {C}_{x,g}/{C}_{{TZ},0}\,\approx \,1/50\). Substituting this relation to Eq. (2), one obtains the dependence of the gelation time on the ratio between NB and TZ groups, \({C}_{{NB},0}/{C}_{{TZ},0}\):$${t}_{g}=\frac{1}{k}\left\{\begin{array}{cc}\frac{1/{C}_{{TZ},0}}{({C}_{{NB},0}/{C}_{{TZ},0}-1)}{ln}\left[\frac{{C}_{{NB},0}/{C}_{{TZ},0}-\beta }{(1-\beta ){C}_{{NB},0}/{C}_{{TZ},0}}\right],&{C}_{{NB},0}\,\ne\, {C}_{{TZ},0}\hfill\\ \frac{\beta }{{C}_{0}(1-\beta )},&{C}_{{NB},0}\,=\,{C}_{{TZ},0}\,=\,{C}_{0}\end{array}\right.$$
(3)
Equation (3) predicts that, for \({C}_{{NB},0}\,\ne \,{C}_{{TZ},0}\), \({t}_{g}\) decreases rapidly with the increase of \({C}_{{NB},0}/{C}_{{TZ},0}\). This theory explains well the experimentally observed gelation time with the fitting parameter \(k\, \approx \,1.8\times {10}^{-2}\,{{{{{{\rm{L}}}}}}}\, {{{{{{{\rm{mol}}}}}}}}^{-1}\,{{{{{{\rm{s}}}}}} }^{-1}\), as shown by the solid line on the left panel of Fig. 2b.For \({C}_{{NB},0}={C}_{{TZ},0}={C}_{0}\), Eq. (3) predicts that the gelation time is inversely proportional to the initial concentration of functional groups. To test this prediction, we synthesize TZ-PAM and NB-PAM polymers with grafting ratios of 2.53% and 2.57%, respectively; this ensures that \({C}_{{NB},0}/{C}_{{TZ},0}\,\approx \,1\). We prepare a series of solution mixtures consisting of equal amount of TZ-PAM and NB-PAM polymers but increase the polymer concentration \(c\) from 1.25 to 15% (w/v). The experimentally measured gelation time decreases from 1352 to 40 s (solid circles in Fig. 2b). This behavior can be well explained by our theory, as shown by the solid line on the right panel of Fig. 2b. Moreover, the fitting parameter, \(k \, \approx \,1.5\times {10}^{-2}\,{{{{{{\rm{L}}}}}}}\, {{{{{{{\rm{mol}}}}}}}}^{-1}\,{{{{{{\rm{s}}}}}}}^{-1}\), agrees reasonably well with that for \({C}_{{NB},0}\,\ne \,{C}_{{TZ},0}\).The gelation theory is based on the physical picture that at the gelation point there is on average one crosslink per polymer. This implies two consequences for the hydrogel stiffness at the gelation point: (i) The stiffness should be constant if the polymer concentration is the same; (ii) the stiffness increases with the polymer concentration. Indeed, these predictions are verified by our experiments. For the hydrogel formulations with the same polymer concentration but various NB/TZ ratios, the shear storage modulus at the gelation time, \({G}^{{\prime} }({t}_{g})\), is nearly a constant of 13 Pa (left panel in Fig. 2c). In contrast, for the hydrogel formulations of the same NB/TZ ratio but various polymer concentrations, \({G}^{{\prime} }({t}_{g})\) increases as the polymer concentration increases (right panel in Fig. 2c). Taken together, our results provide the scientific foundation for prescribed gelation time of the single-network PAM hydrogels in a wide range of concentrations.Stiffness of single-network PAM hydrogelsTo explore the stiffness of single-network PAM hydrogels, we fix TZ-PAM and NB-PAM polymers at a stoichiometrically matched grafting ratio of 2.5% and mix them at equal concentrations from 1.25% to 15% (w/v). We wait for sufficiently long enough time to allow the hydrogel to be completely crosslinked, as exemplified by the saturated \(G^{\prime}\) in Fig. 2a. For all hydrogel formulations, \(G^{\prime}\) is nearly independent of the frequency within the range of 0.1–20 Hz (thick lines, Fig. 2d). Moreover, \(G^{\prime}\) is more than 10 times greater than the loss modulus \(G^{{\prime} {\prime}}\) (thin lines, Fig. 2d). These results indicate that the PAM hydrogels are elastic networks; therefore, one can take \(G^{\prime}\) at the lowest frequency as the equilibrium network shear modulus \(G\).Interestingly, the dependence of hydrogel stiffness on concentration exhibits two distinct regimes. As the polymer concentration increases from 1.25 to 10% (w/v), \(G\) increases rapidly from 385 Pa to 2.1×104 Pa by a power law with an exponent of 2.1 ± 0.2 (dashed orange line, Fig. 2e). Further increasing the polymer concentration results in a negligible increase in stiffness (Fig. 2d, e). This two-regime behavior contradicts conventional understanding that the hydrogel stiffness increases with polymer and crosslinker concentrations.We develop a scaling theory to describe the dependence of hydrogel stiffness on the functionality \({f}_{x}\) and concentration \(c\) of PAM polymers. To form a crosslink, a TZ must meet a NB. Thus, the crosslink concentration is determined by the average distance between a TZ and a NB group. This distance also equals the correlation length, \(\xi\), which is defined as the nearest distance of a monomer from another monomer on the neighboring polymer chains, as indicated by the dashed circles in Fig. 2f. Hence, the size of a network strand, \({a}_{x}\), is approximately the correlation length: \({a}_{x}\,\approx \,\xi\). The shear modulus of the network is about \({k}_{B}T\) per volume pervaded by a network strand:$$G\,\approx \,{k}_{B}T/{{a}_{x}}^{3}\,\approx \,{k}_{B}T/{\xi }^{3}$$
(4)
Because water is a good solvent for PAM, \(\xi\) decreases by \(\xi \,\approx \,{R}_{F}{(c/{c}^{*})}^{-3/4}\), where c* is the polymer overlap concentration, and \({R}_{F}\,{\approx }\,b{N}^{3/5}\) is the end-to-end distance of a single PAM chain in a dilute solution with \(N\) being the number of Kuhn monomers per polymer15. This gives the dependence of the network shear modulus on polymer concentration:$$G\,\approx \,\frac{{k}_{B}T}{{{R}_{F}}^{3}}{\left(\frac{c}{{c}^{*}}\right)}^{9/4}$$
(5)
However, this power law behavior will stop at concentrations greater than \({c}^{*\ast }\), at which the correlation length \(\xi ({c}^{*\ast })\) is about the average distance, \({d}_{x}\), between two neighboring functional groups (TZ or NB) on the same polymer as marked in Fig. 2f. In this case, all functional groups are reacted, and the network shear modulus saturates at:$${G}_{\max }\approx \frac{{k}_{B}T}{{{d}_{x}}^{3}}$$
(6)
In a good solvent,$${d}_{x}\,\approx \,b{\left({f}_{x}^{-1}/{n}_{K}\right)}^{3/5}$$
(7)
Here, \(b=1.6\) nm is the Kuhn length of PAM, \({n}_{K}=\frac{1}{2}{C}_{{\infty }}/{\cos }^{2}\left(\frac{\theta }{2}\right)\,\approx \,6.5\) is the number of chemical monomers per Kuhn monomer with the Flory’s characteristic ratio \({C}_{\infty }=8.5\) and the main C-C chain bond angle \(\theta=68^{\circ}\) 16, and \({f}_{x}^{-1}=40\) is the average number of chemical repeating units within the polymer section of size \({d}_{x}\,\approx \,4.8\) nm. Substituting Eq. (7) into (6), one obtains$${G}_{\max }\,\approx \,\frac{{k}_{B}T}{{\left[b{\left({{f}_{x}}^{-1}/{n}_{K}\right)}^{3/5}\right]}^{3}}\,\approx \,\frac{{k}_{B}T}{{b}^{3}}{\left({n}_{K}{f}_{x}\right)}^{9/5}\propto {{f}_{x}}^{9/5}$$
(8)
Thus, for a fixed \({f}_{x}\), we expect two regimes for stiffness:$$G\,\approx \,\left\{\begin{array}{cc}\frac{{k}_{B}T}{{{R}_{F}}^{3}}{\left(\frac{c}{{c}^{*}}\right)}^{9/4}\propto {(c/{c}^{*})}^{9/4},&{c}^{*} < \,c \, < \,{c}^{*\ast }\\ \frac{{k}_{B}T}{{b}^{3}}{\left({n}_{K}{f}_{x}\right)}^{9/5}\propto {{f}_{x}}^{9/5},&c \, > \, {c}^{*\ast }\hfill\end{array}\right.$$
(9)
At relatively low concentrations, \({c}^{\ast } < \,c \, < \,{c}^{*\ast }\), the network stiffness increases with the polymer concentration by a power of 9/4. In contrast, at high concentration, \(c \, > \,{c}^{*\ast }\), the network stiffness saturates and becomes a constant. The theoretical predictions [see Eq. (9)] explain the experiments remarkably well (Fig. 2e). Note that scaling theory does not include prefactors, which typically have values on the order of unit. For instance, our theory predicts that the maximum hydrogel stiffness is \({G}_{\max }\,\approx \,37\) kPa [Eq. (8)]. Comparing the prediction against the experimentally measured value 22 kPa, one obtains a prefactor of 0.6. Nevertheless, the combination of experiments and theory shows that the TZ/NB-PAM single-network hydrogels allow for prescribed stiffness from ~400 Pa to 22 kPa, sufficient for cell encapsulation and most biomedical applications16,17,18.Stiffness and extensibility of DN hydrogelsTo prepare a DN hydrogel, we use TZ-PAM and NB-PAM, respectively, to prepare two solutions, dissolve the same amount of alginate (Alg) in each of the two solutions, mix the pair of the solutions to initiate the crosslinking of PAM network, and then transfer the partially crosslinked hydrogel to a bath containing 20 mM Ca2+ to incubate for 10 min at RT to completely crosslink both the alginate and PAM networks. We denote a DN hydrogel formulation as PAMxAlgy, where x is the concentration in (w/v)% for TZ-PAM or NB-PAM before mixing, and y is the concentration of alginate. In all formulations, we use TZ-PAM and NB-PAM with stoichiometrically matched grafting ratio of 2.5% at a concentration of 10% (w/v); this formulation results in a single-network PAM hydrogel PAM10 of shear modulus approximately 20 kPa (Fig. 2d), comparable to that of most organs in abdominal space19.Because the alginate network contains carboxylic acid groups and the PAM network contains carboxylic acid and amide groups, and that both the functional groups are sensitive to ions and pH, we equilibrate and completely crosslink hydrogel in cell culture media (DMEM) with 2 mM CaCl2 for 24 h before characterizing its mechanical properties. Indeed, all the hydrogel samples swell in DMEM, as exemplified by a swelling ratio of 7% in size for PAM10Alg2 (Supplementary Fig. 8). The shear modulus of PAM10 hydrogel decreases dramatically from 21 kPa in water to 5.3 kPa in cell culture media. Yet, the hydrogel remains an elastic solid, as evidenced by the nearly frequency-independent \({G}^{\prime}\) as well as small loss factors, \(\tan \delta \equiv {G}^{{\prime} {\prime} }/{G}^{\prime} \,\approx \,0.02\), within the frequency ranging from 0.1 to 20 Hz (solid gray line, Fig. 3a). Similarly, frequency-independent dynamic moduli are observed for 2% (w/v) alginate hydrogel, Alg2, and DN hydrogels (blue and red lines, Fig. 3a). These results show that all hydrogels are non-dissipative, elastic solids.Fig. 3: Stiffness and extensibility of DN polyacrylamide/alginate (PAM/Alg) hydrogels.The grafting ratio of TZ-PAM and NB-PAM is fixed at 2.5% unless otherwise specified. a Dependencies of storage (\(G^{{\prime}}\), solid lines) and loss (\(G^{{\prime} {\prime}}\), dashed lines) moduli of completely crosslinked single-network PAM, single-network alginate (Alg), and double-network PAM/alginate hydrogels on oscillatory shear frequency. The composition of a crosslinked hydrogel is denoted as PAMxAlgy, where x is the concentration in (w/v)% for PAM consisting of equal amount of TZ-PAM and NB-PAM, and y is the concentration of alginate. b Equilibrium shear moduli of single-network PAM, single-network alginate, double-network PAM/alginate hydrogels. c Photos of PAM10, Alg2, and PAM10Alg2 hydrogels under uniaxial tensile tests at a fixed strain rate of 0.02/s. For each formulation, the left, middle, and right panels are, respectively, captured at the beginning of the tensile test, right before breaking, and right after breaking. Samples are 13 mm in length and 2 mm in width of the center part. d Stress–strain behavior of hydrogels. Dashed red line: a PAM10Alg2 double-network hydrogel made from TZ-PAM with a grafting ratio of 2.53% and NB-PAM with a grafting ratio of 0.63%, which is denoted as PAM101/4Alg2. e A two-parameter (\(G\), \({\epsilon }_{f}\)) diagram-of-state that outlines the mechanical properties of all hydrogels formulations in (d). Empty circle corresponds to the hydrogel PAM10Alg2 with mismatched TZ and NB grafting ratios in (d), which is denoted as PAM101/4Alg2.The shear modulus of a DN hydrogel is approximately the sum of the moduli of its constituent singlet-network hydrogels. For instance, Alg2 and PAM10 exhibit shear moduli of 4.5 kPa and 5.3 kPa, respectively, resulting in a combined modulus of 9.8 kPa, which closely matches the value of 11 kPa for PAM10Alg2. This observation holds true for other DN hydrogel formulations as well. When the concentration of alginate increases from 1% to 3% (w/v), the stiffness of single-network alginate hydrogels increases linearly from 1.2 kPa to 6.6 kPa (dashed bars, Fig. 3b). The corresponding DN hydrogels show a similar increase in stiffness from 8.2 kPa to 14 kPa with a nearly the same magnitude (solid bars, Fig. 3b).Next, we perform uniaxial tensile tests to quantify the extensibility of DN hydrogels. To do so, we take a fully crosslinked hydrogel film measuring approximately 1 mm in thickness and cut it into a dog-bone shape sample. We then subject the sample to a fixed strain rate of 0.02/s while capturing the process using a camera, as visualized by the photos in Fig. 3c and by Supplementary Movie 1. For all hydrogels, we find that the stress increases almost linearly with tensile strain until the hydrogels reach their breaking point. However, we note that the single-network Alg2 hydrogel is often too brittle for reliable measurements. These findings further confirm that the hydrogels are nearly pure elastic networks with little energy dissipation.The DN PAM10Alg2 hydrogel exhibits significantly improved extensibility compared to the single-network PAM10 hydrogel. While PAM10 has a tensile breaking strain \({\epsilon }_{f}\) of 0.36, the PAM10Alg2 hydrogel demonstrates approximately three times greater extensibility with \({\epsilon }_{f}\) of 1.0, as shown in Fig. 3c, d and by Supplementary Movie 1. Remarkably, although increasing the alginate concentration stiffens the corresponding DN hydrogels, \({\epsilon }_{f}\) remains nearly the same at 1 (solid lines in Fig. 3d and filled circles in Fig. 3e).We attribute the observed enhancement in network extensibility to the characteristics of DN hydrogels. Upon deformation, the weak alginate network fractures first; this process effectively prevents localized and amplified stress along PAM network strands, thereby improving network extensibility. However, the network extensibility cannot exceed that afforded by the network strand of the PAM network. In theory, the maximum elongation at break, \({\lambda }_{\max }^{T}\), for a network strand equals the ratio between its initial size, \({a}_{x}\), and contour length, \({L}_{\max }\):$${\lambda }_{\max }^{T}=\frac{{L}_{\max }}{{a}_{x}}$$
(10)
However, typical single-network hydrogels, which often have a wide distribution in network strand size, cannot reach this maximum extensibility. When subjected to deformation, the relatively short network strands would break first, resulting in extensibility lower than the prediction based on average network size \({a}_{x}\). By contrast, in a DN hydrogel, the weak network can undergo fracture to prevent localized, amplified stress near network defects or along network strands, and therefore, avoid premature failure of the strong network. Consequently, a DN hydrogel may reach the theoretical extensibility \({\lambda }_{\max }^{T}\)20,21.We develop a scaling theory to describe the dependence of network theoretical extensibility \({\lambda }_{\max }^{T}\) on network stiffness \(G\). In a DN PAM/alginate hydrogel, the PAM network has on average \({N}_{x}\) Kuhn monomers per network strand. In a good solvent, the end-to-end distance, or the size of the network strand, is a self-avoiding random walk of Kuhn monomers with size of \(b\):$${a}_{x}\,\approx \,{b{N}_{x}}^{3/5}$$
(11)
Recall Eq. (10), the theoretical extensibility of the network is$${\lambda }_{\max }^{T}\,\approx \,{{N}_{x}}^{2/5}$$
(12)
For an unentangled network, its shear modulus \(G\) is about \({k}_{B}T\) per volume pervaded by the network strand,$$G\,\approx \,\frac{{k}_{B}T}{{a}_{x}^{3}}\,\approx \,\frac{{k}_{B}T}{{b}^{3}{N}_{x}^{9/5}}$$
(13)
Thus, one can correlate \({\lambda }_{\max }^{T}\) to the network stiffness \(G\),$${\lambda }_{\max }^{T}\,\approx \,{({k}_{B}T/G)}^{2/9}{b}^{-2/3}\propto {G}^{-2/9}$$
(14)
Equation (14) predicts that extensibility of the DN hydrogel increases with the decrease of PAM hydrogel stiffness.To test this prediction, we synthesize another DN hydrogel, PAM101/4 Alg2, in which the TZ-PAM has a grafting ratio of 2.53%, while the NB-PAM has a mismatched grafting ratio of 0.63%. This formulation results in a DN hydrogel with lower network shear modulus, \(G\,\approx \,7.4\) kPa (Supplementary Fig. 9) and higher maximum extensibility, \({\lambda }_{\max,1/4}\,\approx \,2.3\) (dashed line in Fig. 3d and empty circle in Fig. 3e). Considering that the ionically crosslinked alginate network has negligible contribution to \({\lambda }_{\max }^{T}\) and the equilibrium shear modulus \({G}_{1/4}\), \({G}_{1/1}\), and \({G}_{{alg}}\) are respectively 7.4, 11.5,and 4.5 kPa (Fig. 3b, e), the ratio of the extensibility of the softer DN hydrogel with NB/TZ ratio 1/4 to that of the stiffer one with NB/TZ ratio 1/1 is predicted to be \({\lambda }_{\max,1/4}^{T}/{\lambda }_{\max,1/1}^{T}\,\approx \,{[({G}_{1/4}-{G}_{{alg}})/({G}_{1/1}-{G}_{{alg}})]}^{-2/9}\approx \,1.2\). Remarkably, this value agrees well with the experimentally measured ratio 1.2. These results collectively show that the extensibility of the DN hydrogel is determined by the PAM network.Notably, compared to the classical DN PAM/alginate hydrogel of the same composition, which exhibits a remarkable extensibility with \({\epsilon }_{f}\,\approx \,20\)22, our DN hydrogel is much less stretchable. This significant difference in absolute extensibility is likely due to the difference in network topology. In the classical DN hydrogel, the PAM network is formed through in situ free radical polymerization of acrylamide monomers in an alginate solution. This polymerization is an uncontrolled reaction, resulting in a wide distribution in network strand size and leading to relatively large extensibility of a single-network PAM hydrogel with \({\epsilon }_{f}\,\approx \,6.5\). In contrast, in our DN hydrogels, the PAM network is formed through the click reaction between TZ-PAM and NB-PAM precursor polymers, which have a relatively low and fixed molecular weight. Furthermore, the reactive groups further segment the linear PAM polymers into small sections, thereby further reducing the network’s extensibility.DASP printing of DN hydrogelsDASP printing requires highly viscous and shear-thinning bio-inks, such that during printing, an extruded droplet can grow uniformly in the yield-stress fluid supporting matrix9. We conduct shear tests to measure the viscosity of the DN PAM10Alg2 bio-ink and its constituent PAM10 and Alg2 solutions. The PAM10 solution is nearly a Newtonian fluid with a low viscosity of 0.2 Pa·s, which is nearly independent of shear rate ranging from 0.1 to 1000 s−1, as shown by the squares in Fig. 4a. In contrast, the Alg2 solution is a shear-thinning fluid, with the apparent viscosity decreasing with the shear rate by a power of −0.63, as shown by the triangles in Fig. 4a. However, the magnitude of the viscosity, 2.6 Pa·s, remains much lower than ~20 Pa·s required for DASP printing. Consequently, both constituent bio-inks are unsuitable for DASP printing. Remarkably, the viscosity of the DN bio-ink significantly increases to 17.1 Pa·s while maintaining shear-thinning behavior, as shown by the solid circles in Fig. 4a. These properties make the DN bio-ink PAM10Alg2 suitable for DASP printing9. Moreover, the crosslinked DN hydrogel has a significantly improved extensibility compared to its single-network constituents and a stiffness comparable to that of the organs in the abnormal cavity. Thus, we use PAM10Alg2 as the DN bio-ink for the subsequent DASP printing.Fig. 4: 3D printing of DN hydrogels.a The dependence of viscosity of the PAM (10%, w/v), Alg (2%, w/v), and the corresponding DN bio-ink, PAM10Alg2, on the shear rate, \(\dot{\gamma }\), at RT. b A dual-inlet print nozzle with a static mixer chamber. (i) A 3D rendering of the static mixer chamber. (ii) Photos of the nozzle with a pair of DN PAM10Alg2 bio-inks flowing through the mixer chamber. The two bio-inks contain red and green microparticles, respectively, with a diameter of 50 μm. The left, middle, and right panels of (ii) are sequential photos when only the red channel flows, both channels flow, and only the green channel flows, respectively. Scale bar, 1 mm. c A DASP printed hollow sphere consisting of 42 interconnected yet distinguishable DN PAM10Alg2 hydrogel particles. Panels (i), (ii), and (iii) display the front view, computer-rendered model, and the hollow sphere that is cut into two pieces. Scale bar for panel (iii) is 1 mm. d A gyroid structure created by conventional printing that uses DN PAM10Alg2 hydrogel filaments as building blocks. Panels (i), (ii), and (iii) respectively depict the 45° view, computer-rendered 3D model, and side view of the gyroid. Scale bar for panel (iii) is 5 mm. e Photos of the gyroid structure from (d) undergoing cyclic compression testing with a compression strain rate of 0.005/s. Panels (i), (ii), (iii), and (iv) are captured at 0%, 30%, 60% compression strain, and full release in the first compression cycle. Scale bar, 5 mm. f Mechanical behavior of the gyroid structures during cyclic compression. The gyroids are made of DN PAM10Alg2 (solid lines) and single-network Alg2 (dashed lines) hydrogels. Insert: a zoom-in view of the compression curves.For DASP printing, the bio-inks are required to be extruded through the nozzle in a liquid state and rapidly crosslink once being deposited. Due to the rapid crosslinking of DN bio-inks upon mixing (Fig. 2a), premixing and loading the DN bio-inks into a single syringe, as performed in DASP 1.0, would clog the print nozzle and syringe. To avoid the clogging, we separately load the TZ-PAM and the NB-PAM in two syringes and then mix them on-demand during the printing. To do so, we engineer a dual-inlet print nozzle with an inner static mixer chamber (Supplementary Fig. 10). The design of the mixer chamber utilizes a structure called Quadro™ Square, as depicted by a computer-rendered 3D model in Fig. 4b, i. We use stereolithography printing to fabricate this nozzle, which has an outer diameter of 2.7 mm. The inner channel of the nozzle has a narrowest dimension of 425 μm. These dimensions are much larger than the size of typical cells, ~30 μm23, and even some large cell aggregates such as islets, ~200 μm24. This design avoids mechanical shear-induced damage to the cells during extrusion.To evaluate the mixing capability of the dual-inlet print nozzle, we use a camera to monitor the mixing process in real-time. To visualize the mixing, we incorporate red and green mica microparticles with the diameter of 50 μm into the respective bio-inks. The pair of bio-inks is extruded through the two inlets of the nozzle at a flow rate of 0.15 μL/s, which is the same as that used in DASP printing. Initially, we turn on only the red channel, allowing the mixer chamber to be filled with red particles (left, Fig. 4b, ii). Subsequently, we switch on the green channel, and within 120 s, the mixer chamber transitions to a greyish purple color because of thorough mixing (middle, Fig. 4b, ii). During this mixing process, approximately 36 μL of bio-ink passes through the chamber, which is around the volume of the chamber itself. Finally, we turn off the red channel, and within another 240 s, the chamber completely changes to a green color (right, Fig. 4b, iii). The smooth transition of colors in the mixer chamber demonstrates the nozzle’s capability to mix the viscoelastic bio-inks homogeneously (Supplementary Movie 2).To showcase the capabilities of DASP 2.0, we seek to fabricate a structure reminiscent of a raspberry: A hollow sphere with a shell composed of a single layer of hydrogel particles. Achieving this goal perhaps represents one of the most formidable tasks in voxelated bioprinting, as it demands precise generation, deposition, and assembly of individual particles in 3D space. Furthermore, it requires a robust connection between neighboring hydrogel particles to ensure the mechanical integrity of the structure.Building on our previously established knowledge and leveraging the newly developed extrusion module, the printed droplets in the supporting matrix undergo isotropic swelling by 14% in diameter (Supplementary Fig. 11), enabling them to coalesce into a hollow sphere consisting of only one layer of interconnected yet distinguishable DN hydrogel particles. This sphere has a diameter of 7 mm and comprises 42 hydrogel particles, each approximately 1 mm in diameter, as shown by a representative photo in Fig. 4c, i and illustrated by the 3D rendering in Fig. 4c, ii. All the particles are interconnected yet distinguishable, resulting in a mechanically robust free-standing hollow sphere that can be easily manipulated without additional protection (Supplementary Movie 3). Furthermore, we confirm the hollow nature of the sphere by cutting it into two pieces, revealing a single layer of hydrogel particles (Fig. 4c, iii; Supplementary Movie 4). Compared with the simple cubic lattice structure printed in our previous work8, fabricating such a hollow sphere with high surface curvature represents a greater challenge. This is primarily due to the risk of the structure collapsing if the hydrogel itself is mechanically brittle. Therefore, the successful fabrication of this hollow sphere using DASP 2.0 represents a significant advancement in voxelated bioprinting.In a DASP-printed structure, the bridges between neighboring droplets may be weaker than the bulk hydrogel, which would impair the mechanical robustness of the structure. To this end, we quantify the extensibility of a DASP-printed 1D filament that comprises a sequence of interconnected droplets (Supplementary Fig. 12). As expected, the 1D filament has a tensile breaking strain \({\epsilon }_{f}\) of 0.7, which is slightly lower than that of bulk hydrogel of 1.0 (Fig. 3c, d). Nevertheless, the breaking strain of the 1D filament remains much larger than the constituent single-network hydrogels (\({\epsilon }_{f}\approx 0.36\)). These results further support that DASP enables the fabrication of integrated structures consisting of interconnected droplets.DASP 2.0 can be easily converted to conventional one-dimensional (1D) filament-based bioprinting, as demonstrated by the creation of a gyroid cubic structure (Fig. 4d). Unlike other mesh-like structures such as waffles or honeycombs in which the channels are separated by walls, the gyroid is a highly porous structure consisting of an interconnected network of channels and voids (Fig. 4d, ii). To 3D print such a complex geometric pattern, each layer must slightly protrude beyond the previous layer, forming the so-called hangovers. This poses a challenge for 3D bioprinting, as bio-inks often lack the mechanical strength to support the subsequent hangover layer. However, this caveat can be circumvented using our DN bio-ink, which not only is mechanically strong but also possesses relatively fast crosslinking kinetics. Indeed, with DASP 2.0, we successfully print our DN bio-ink into a gyroid with high fidelity, as confirmed by the 3×4 array of tunnel-like holes with a diameter of 1.5 mm on the side view (Fig. 4d, i & iii), precisely capturing the intended design (Fig. 4d, ii; Supplementary Movie 5).To assess the mechanical properties of the DN gyroid, we perform a cyclic compression test at a strain rate of 0.005/s. The DN gyroid can sustain a large compression strain of 60% without fracturing (Fig. 4e, i-iii). Upon stress release, it completely recovers its original height, as visually depicted in Fig. 4e, iv and Supplementary Movie 6. The stress-strain profiles during loading and unloading almost overlap (solid lines, Fig. 4f). Quantitatively, the energy dissipation efficiency, defined as the ratio between the integrated area in the hysteresis loop and that under the compression curve, is relatively small with a value of 27.5%. This elastic, non-dissipative behavior persists over eight cycles of cyclic compression, as evidenced by the nearly perfectly overlaying of stress-strain profiles (solid lines, Fig. 4f). In contrast, the gyroid made of pure alginate exhibits plastic, dissipative behavior with the energy dissipation efficiency remarkably high of 80.2% at the first cycle. After eight compression cycles, the height of the gyroid decreases to 74% of its original value, as indicated by the increase of onset strain above which the stress becomes positive (dashed line, inset of Fig. 4f). Moreover, as the number of compression cycles increase to eight, the stress at 60% strain decreases by approximately 30% from 8.6 kPa to 6.0 kPa. These results demonstrate that DASP printing DN bio-inks enables the fabrication of highly deformable, elastic, non-dissipative, and multiscale porous scaffolds.Cytocompatibility and in vivo stability of DASP printed DN scaffoldsCompared to traditional scaffolds based on bulky hydrogels, the DASP-printed DN scaffolds offer a promising platform for cell-based therapy, which often involves encapsulating therapeutic cells for transplantation. The multiscale porosity has previously been shown to facilitate efficient nutrient transport8, while the mechanical robustness is expected to allow the scaffolds to withstand constant mechanical perturbation resulting from the movement of the transplant recipient. To demonstrate this potential, we begin by investigating the cytocompatibility of both the printing process and the bio-ink by encapsulating cells into each hydrogel droplet. To do so, we modify the print nozzle by adding an extra channel, creating a three-channel nozzle (Supplementary Figs. 10b, c). Similar to the original dual-inlet print nozzle, two channels are used to load the DN ink. However, the third channel is dedicated to loading cells, as illustrated in Fig. 5a, i. This design ensures that the cells are not loaded directly to DN ink, as this could change the rheological properties of the ink. Such alterations can be significant and unpredictable at high cell density, rendering the ink unsuitable for DASP printing. Additionally, following our previously developed method25, we incorporate 15% (w/v) dextran and 15% (w/v) poly(ethylene glycol) (PEG) polymers into the cell suspension and the DN ink, respectively. Compared to the DN ink, these polymer solutions are of relatively low viscosity and do not impact the printing process. This method results in an aqueous two-phase system, in which both the dextran and PEG solutions are aqueous and cytocompatible yet have a small interfacial tension that is sufficient to lead to phase separation26. During the extrusion process, the DN ink encounters the dextran phase in the static mixer chamber (Fig. 5a, i). After extrusion, the dextran phase is expected to phase separate from DN ink, forming an island-sea-like morphology, as illustrated in Fig. 5a, ii. Indeed, this microstructure is confirmed by visualizing the distribution of dextran phase and cell-mimicking polystyrene microspheres with a diameter of 20 μm, respectively shown by a representative fluorescence confocal microscopy 3D image in Supplementary Fig. 13 and a bright field microscopy image in Fig. 5a, iii (see Methods). These results demonstrate that, together with the aqueous two-phase system, the three-channel nozzle allows for transforming the DN ink into a sponge-like structure with a solid hydrogel filled with cell-laden liquid voids.Fig. 5: Cytocompatibility and in vivo stability of DASP printed DN polyacrylamide/alginate (PAM/Alg) scaffolds.The single-network and DN scaffolds are respectively made of Alg4 and PAM10Alg2 hydrogels. a A strategy that utilizes PEG/dextran aqueous two-phase system to encapsulate cells during DASP printing. (i) A 3D rendering of the print nozzle for cell encapsulation. The print nozzle consists of three channels: one channel loaded with cell suspension in dextran (green flow) and two channels respectively loaded with the pair of bio-inks mixed with PEG (red and blue flow). During the extrusion, the flows of bio-ink first encounter with the immiscible flow of the cell suspension and subsequently mix in the static mixer chamber. (ii) A schematic of the extruded bio-ink in which the cell suspension phase (green) is separated from the bio-ink phase (red). (iii) A representative microscopy image showing the distribution of cell-mimicking polystyrene microparticles in the extruded bio-ink. Scale bar, 50 μm. b A representative fluorescence confocal microscopy image from live/dead assay of DASP encapsulated Beta-TC-6 cells. Scale bar, 100 μm. c Viability of naked and DASP encapsulated Beta-TC-6 cells up to 3 days. d Glucose stimulation index of naked and DASP Beta-TC-6 cells in 3 days. Results in (c) and (d) are shown as mean ± standard deviation (S.D.) with sample size n = 5 wells for naked Beta-TC-6 and n = 5 scaffolds for DASP Beta-TC-6 in PAM/Alg. e An animal model for testing the mechanical robustness of DASP printed DN scaffolds in vivo. (i) A schematic of transplanting a DASP printed 5×5×4 lattice scaffold into the abdominal cavity of immunocompetent C57BL/6 mice. (ii) A representative photograph of the surgery. f Representative photos of the scaffolds made of pure alginate hydrogel (i) before the transplantation and (ii) after being retrieved. Scale bars, 1 mm. g Representative photographs of the scaffolds made of DN hydrogel (i) before the transplantation, (ii) after being retrieved, (iii) located in the abdominal cavity (dashed circles), and (iv) under digital camera after being retrieved. Scale bars, 1 mm. h Mechanical behavior of scaffolds retrieved at different time points under compression.By replacing the cell-mimicking microspheres with Beta-TC-6 cells, we print a 5×5×4 DN lattice (Supplementary Fig. 14). Beta-TC-6 cells are a pancreatic beta cell line that retains the function of glucose-stimulated insulin secretion, allowing us to test both the cell viability and the transport properties of the multiscale porous scaffold. The live/dead assay reveals that the viability of Beta-TC-6 cells immediately after printing is 61.3 ± 8.0%, and it remains above 60% for the subsequent three days (Fig. 5b, c). This viability is consistent with the values 40-80% reported for extrusion-based bioprinting27, but it is considerably lower compared to DASP 1.0, which uses pure alginate as the hydrogel and exhibits a viability of 90%8. The lower viability can be attributed to the cytotoxicity of free polyacrylamide polymers present in the DN ink before complete crosslinking. Although polyacrylamide-based hydrogels are generally considered as non-toxic and extensively used in biomedical research, our study reveals that free polyacrylamide polymers exhibit some degree of cytotoxicity. Indeed, live/dead assay conducted on cells directly mixed with pure polyacrylamide polymers with the same concentration as that used for the DN bio-ink shows a cell viability less than 1% (Supplementary Fig. 15). Nevertheless, in DASP 2.0 more than half of the cells survive and maintain their viability for three days, which is a typical waiting time before transplantation in clinical settings.Next, we test the insulin release of the lattice scaffolds using static glucose-stimulated insulin secretion (GSIS) measurement. Specifically, we expose the Beta-TC-6 cell-encapsulated scaffolds to low and high glucose media and measure the amount of released insulin. The release index is defined as the ratio between insulin release amount in high glucose media over that in low glucose media. On the first day, the lattice scaffolds exhibit a relatively low release index with a value of 1.39 ± 0.56. Yet, this value is comparable to that of the naked counterpart, which exhibits a release index of 1.31 ± 0.42 (Fig. 5d). This is likely because that the cells are in a recovery phase after detachment from the culture flask. In the subsequent days, the release index of the naked cells increases significantly, reaching a value of 6.78 ± 1.03 on day 3, which is about twice the value of 2.98 ± 0.83 observed for the lattice scaffolds (Fig. 5d). This difference is consistent with our previous studies, in which we found that the release index of islets encapsulated 3×3×3 lattice scaffolds created using DASP 1.0 is about half that of the naked counterpart. Nonetheless, the release index of DASP Beta-TC-6 cells on day 3 is much higher than one, which exceeds the standard threshold for clinical transplantation. Taken together, our results show that DASP 2.0 enables the encapsulation of living cells while maintaining reasonable cell viability and function.To further confirm the suitability of the DN scaffolds for clinical transplantation, we perform in vivo studies to evaluate their mechanical robustness. We transplant the 5×5×4 lattice scaffolds into the abdominal cavity of immunocompetent C57BL/6 mice, as respectively illustrated and photographed in Fig. 5e, i and ii. This in vivo environment is characterized by vigorous mechanical perturbation resulting from intestinal peristalsis and body movement. Two bio-ink formulations are used to prepare the scaffolds: (i) 4% (w/v) alginate, our previous prototype formulation for DASP 1.0, denoted as Alg4, and (ii) DN bio-ink for DASP 2.0, a formulation explored for cell encapsulation, denoted as PAM10Alg2 (Fig. 5a). Both Alg4 and PAM10Alg2 scaffolds exhibit similar stiffness, with respective apparent compression moduli of 5.6 and 5.8 kPa (dashed red line and dark line, Fig. 5h). However, a significant difference in in vivo stability is observed for these two scaffolds. Alg4 scaffolds fragment only 7 days after the transplantation and can be hardly retrieved, as shown in Fig. 5f. This phenomenon can be attributed to the fragility of the alginate scaffold, which fractures under a relatively small strain of 37% during compression test (dashed red line, Fig. 5h), and limited stability of ionic crosslinks in physiological conditions28. In contrast, the PAM10Alg2 scaffolds retain their integrity without observable deformation for 120 days after the transplantation, as shown by Fig. 5g, i and ii. As observed in different mice, the PAM10Alg2 scaffolds are widely distributed in the abdominal cavity, in contact with the stomach, intestine (Fig. 5g, iii), bladder, liver (Supplementary Fig. 16a), and sometimes slightly adhering to fat tissue, yet all scaffolds could be successfully retrieved (Supplementary Fig. 16b). Moreover, the retrieved scaffolds maintain the pre-transplant lattice structure in which individual particles are interconnected yet distinguishable (Fig. 5g, iv).The integrity of the retrieved DN scaffolds is further confirmed by their mechanical robustness. For scaffolds retrieved 30 days after transplantation, their yield compression stress decreases from 39.3 kPa to 23.8 kPa compared to its pre-transplant value; however, the yield compression strain remains remarkably high of approximately 60% (solid blue line, Fig. 5h). Moreover, the stress-strain curve nearly perfectly overlaps with that of the pre-transplant one, and this behavior persists for scaffolds retrieved after longer periods of transplantation for 60 and 120 days (purple and green lines, Fig. 5h). Taken together, our results show that the DASP 2.0 printed DN scaffolds possess sufficient mechanically robustness to maintain long-term in vivo stability.Universality of DN (A + B)/C hydrogels as bio-inks for DASP 2.0We note that the cytocompatibility of the PAM/alginate DN bio-ink is relatively low, with cell viability ~60% (Fig. 5c), largely because free PAM-based polymers are not completely cytocompatible (Supplementary Fig. 15). Yet, the concept of (A + B)/C DN hydrogels based on TZ-NB click chemistry is expected to be applicable to other biopolymers, where A, B, and C respectively represent TZ-polymer, NB-polymer, and alginate. To test this, we replace the PAM polymer by hyaluronic acid (HA), a natural polysaccharide widely used in biomaterials development. Following the same chemistry as for the PAM (Fig. 1b), we functionalize a pair of HA polymers, respectively, with TZ and NB at nearly the same grafting ratio of 27%, corresponding to approximately 250 TZ or NB groups per polymer (Supplementary Figs. 6, 7). Upon mixing, the pair of TZ-HA and NB-HA crosslink to form a network via additive-free click reaction, as illustrated in Fig. 6a.Fig. 6: Mechanical properties, printability, and biocompatibility of a DN hyaluronic acid/alginate (HA/Alg) bio-ink.a HA is functionalized with either norbornene (NB) or tetrazine (TZ) groups. Upon mixing, TZ-HA and NB-HA are crosslinked via the click reaction between TZ and NB groups, following the same crosslinking mechanism as that for creating a PAM network (Fig. 1b). The grafting ratio of TZ-HA and NB-HA are both fixed at 27%. b Dependencies of storage (\(G^{\prime}\), solid lines) and loss (\(G^{{\prime} {\prime}}\), dashed lines) moduli of completely crosslinked single-network HA and DN HA/alginate hydrogels on oscillatory shear frequency. Single-network hydrogel (HA): TZ-HA (2%, w/v) + NB-HA (4%, w/v). The DN hydrogel (HA/alginate) is prepared by mixing a pair of bio-inks, TZ-HA (2%, w/v) / alginate (2%, w/v) + NB-HA (4%, w/v) / alginate (2%, w/v). c Representative photos of the DN HA/alginate hydrogel under uniaxial tensile test at a fixed strain rate of 0.02/s. The left, middle, and right panels, respectively, are captured at the beginning of the tensile test, right before fracture, and right after fracture. Samples are 13 mm in length and 2 mm in width of the center part. d Stress-strain behavior of the single-network HA and DN HA/alginate hydrogels in (b). e A DASP printed 5×5×4 lattice scaffold consisting of 100 interconnected yet distinguishable DN HA/alginate hydrogel particles. Upper panel: top view; lower panel: side view. Scale bar, 1 mm. f Cytocompatibility of the DN HA/alginate bio-ink. (i) A pair of HA/alginate bio-inks pre-loaded with cells are printed using a two-channel print nozzle. (ii) A representative fluorescence confocal microscopy image from live/dead assay of DASP encapsulated Beta-TC-6 cells within the DN HA/alginate hydrogel. Scale bar, 100 μm. (iii) Viability of naked and DASP encapsulated Beta-TC-6 cells within the DN HA/alginate hydrogel up to 3 days. Results are shown as mean ± standard deviation (S.D.) with sample size n = 5 wells for naked Beta-TC-6 and n = 5 scaffolds for DASP Beta-TC-6 in HA/Alginate. g Representative photographs of the scaffolds made of DN HA/alginate hydrogel (i) before the transplantation, (ii) after being retrieved at 30 days. Scale bar, 1 mm.Compared to the single-network hydrogels, the improvement in the mechanical properties of the HA/alginate DN hydrogel is comparable to that observed for the PAM/alginate formulation. For instance, the shear modulus of a single-network HA is about 500 Pa (solid gray line, Fig. 6b). By contrast, for the DN HA/alginate hydrogel, G is about 3050 Pa, approximately the sum of HA and alginate single-network hydrogels (red solid line, Fig. 6b). Moreover, unlike the single-network HA hydrogel that is brittle with a tensile breaking strain of 0.6, the DN HA/alginate hydrogel is significantly more stretchable with \({\epsilon }_{f}=1.02\), as shown in Fig. 6c, d and Supplementary Movie 7.As expected, the HA/alginate is suitable for DASP printing, as demonstrated by the success of creating a 5×5×4 lattice scaffold (Fig. 6e). To test the cytocompatibility of the HA/alginate hydrogel, we directly mix Beta-TC-6 cells with the bio-ink at a cell density of 5 million/mL. The live/dead assay reveals that the viability of HA/alginate encapsulated Beta-TC-6 cells remains above 80% for three days after the printing (Fig. 6f). This survival rate is significantly higher than 60% observed in PAM10Alg2, and importantly, meets the standards for most applications27. Finally, we transplant the DASP printed 5×5×4 lattice scaffolds made of HA/alginate into the mouse abdominal cavity. The integrity of the retrieved scaffolds is confirmed 30 days after transplantation (Fig. 6g), indicating that HA/alginate DN scaffolds are sufficiently robust to maintain long-term in vivo stability. Taken together, these results demonstrate that the concept of (A + B)/C hydrogels provides a universal strategy for the development of modular DN bio-inks for DASP printing.
Voxelated bioprinting of modular double-network bio-ink droplets
