The diffusion of three model proteins, namely lysozyme, BSA and Ig, were chosen for this study for several reasons: 1) their physical properties are well characterized (Peters 1996
), 2) they all have rs
smaller than the calculated ξ of the hydrogels, and 3) they cover a large range of sizes. If the protein release is to be controlled by the degradation of the polymer, the comparison of protein size and ξ is very important; for example, in cases where initial gel ξ is much larger than the protein to be released, the protein diffuses out without significant influence from gel degradation. The release is influenced primarily by the hydrogel degradation when the protein size is similar or larger than the initial gel ξ. Thus, control of ξ is very important in the design of controlled release devices and scaffolds for tissue engineering applications, and for this reason, parameters that influence degradation, ξ and diffusivity were investigated in this study. All hydrogels in the study were hydrolytically degradable; hence the changes in ξ and diffusivity were monitored over time. A non-degradable control was also characterized for comparison.
The ξ and protein release were controlled by three major strategies: the molecular weight of the PEG polymer, the number of methylene groups between the ester and the thiol of the cross-linker, and the polymer density. We confirmed the findings of similar studies (Bell and Peppas 1996
; Cruise et al. 1998
; Lu and Anseth 2000
) that all the chosen strategies were efficient in controlling ξ as increase in molecular weight of the PEG polymer, decrease in the number of methylene groups between the thiol and the ester of the cross-linker, and decrease in polymer density all lead to an overall increase in ξ as well as increase in the rate of change of ξ (). However, the number of methylene groups between the ester and the thiol of the cross-linker had the largest effect on ξ. To explain this result we first must note that the number of methylene groups had a profound effect on hydrogel degradation (gels with only 1 methylene group degraded in 16 h as opposed to gels with 2 methylene groups that degraded in 6 d). As the cross-linkers used in forming the hydrogel were degradable via an ester bond incorporated in the polymer backbone, the presence of water causes the ester bond to hydrolyze, resulting in breakage of a cross-link and increase in ξ. As cross-linkers with esters in close proximity to thiols hydrolyze more quickly than those with greater separation between the ester and thiol (Schoenmakers et al. 2004
; Rydholm et al. 2007
; Zustiak and Leach 2010
), a more rapid change in ξ would be observed for gels made with fewer methylenes between the ester and thiol. In all cases, increase in ξ enhanced the diffusion of proteins not only initially but also further upon degradation.
Based on degradation stage and gel type, ξ varied from ~13 to ~35 nm. All proteins in this study had a smaller rs than the calculated ξ; therefore we assumed that all of the proteins had unobstructed diffusion in the gel. Even so, protein size proved to be an important factor in controlling diffusivity. The protein release rate was inversely proportional to protein size and was more sensitive to physical size rather than other protein properties such as charge, which was an expected result as PEG is inert (). We observed that the smallest protein, lysozyme, with a diameter that was several orders of magnitude smaller than the mesh size of the PEG gel, was completely released in less than 24 h. Since the gel degradation was negligible in this time frame (), the release of lysozyme was guided by simple diffusion. On the other hand, the largest protein, Ig, with a diameter less than an order of magnitude smaller than the initial mesh size of the gel, exhibited a prolonged release over several days indicating that the release was aided by gel degradation as well as diffusion. Release of this protein was steady and did not display a burst at complete gel degradation signifying the potential use of the presented gels for sustained release of large protein molecules in drug delivery or tissue engineering applications.
Note that shows that the lysozyme release profiles start with a burst at early times (<200 min). Since the size of lysozyme was relatively small compared to ξ, lysozyme diffused freely and a higher initial protein concentration gradient may have contributed to the burst effect. Other factors that may have added to this burst include: 1) lysozyme molecules that were near the solvent/gel interface and escaped rapidly in the surrounding solution, and 2) faster release of lysozyme through larger pores of the gel compared to slower release from pores obstructed by polymer entanglements.
Based on we observed that the application of the commonly used Fickian model (Eq. 2
) resulted in De
significantly different from Do
(~94% for Ig). The model assumes: a) unobstructed diffusion of small solutes inside the gel, b) infinite dilution of the solute in the supernatant, and c) diffusion inside the gel based solely on Brownian motion. Nevertheless, Eq. 2
is a good approximation even when not all the assumptions are met and has been widely used for calculating diffusivities in non-ideal gel systems (Leach and Schmidt 2005
; Koutsopoulos et al. 2009
). The deviation of experimental results from the model, however, would increase as the size of the protein approaches ξ, as was the case of Ig which has a hydrodynamic diameter of 9.4 nm (~ 70% of the average ξ, 14 nm). Even small proteins such as lysozyme (3.2 nm diameter) were obstructed by the polymer chains which resulted in ~50% decrease in De
. Furthermore, even though we assumed a homogeneous matrix, it is possible that there were some non-ideal physical entanglements in the PEG chains which additionally decreased ξ or non-specific protein/polymer interactions that contributed to decreased De
(Zustiak et al. 2010
). Also note that the initial concentration of all proteins inside the gels was 2% w/v; hence, protein crowding may have affected diffusion by slowing molecular motion. The rationale for using a relatively high concentration of protein was to resemble application wherein a delivery vehicle or 3D scaffold would carry a high load of bioactive molecules.
Next, we aimed to identify simple and effective ways to control protein release. We saw that at initial time points (< 16 h), degradation of the gel was negligible, and protein diffusion as represented by Mi
, was correlated to ξ (Peppas et al. 1999
; Tsunomori and Ushiki 1999
). This relationship can be explained by Eq. 2
in conjunction with Eq. 7
which applies to neutral uncharged polymers such as PEG (Lustig and Peppas 1988
This equation suggests that protein diffusivity, as calculated from Mi
as per Eq.2
, is dependent on ξ and will decrease with increase in solute size. As expected, based on the results from and the dependence of protein diffusion on ξ, an increase in BSA diffusivity was associated with increased polymer molecular weight, decreased number of methylene groups between the thiol and the ester of the cross-linker, and decreased polymer density (, and ). Similar trends have been noted previously for various degradable PEG polymers (West and Hubbel 1995
; Jeong et al. 2000
; Schoenmakers et al. 2004
). In all of the above cases we saw that the diffusion data for Mi
< 0.6 was well explained by the Fickian diffusion model (Eq. 2
). This finding suggested that initial protein release was guided mainly by diffusion. However, in the case of the fast-degrading hydrogels made with PEG-SH 1 3.4, we observed a ~3-fold increase in BSA De
as compared to the slow-degrading gels made with PEG-SH 2 3.4. Therefore, we speculate that bulk degradation of the hydrogel (due to the hydrophilic nature of the PEG polymer (Sawhney et al. 1993
)) and subsequent increases in gel water content and ξ were partially responsible for this result.
It has been shown that encapsulating protein in hydrogels prior to gelation may affect protein stability and structure (Morlock et al. 1997
; Eggers and Valentine 2001
; Wetering et al. 2005
). The effects include but are not limited to various types of degradation such as denaturation, aggregation, hydrolysis, and reaction with the cross-linkers which could lead to polymer-protein adducts. Such effects have the potential to decrease protein activity after release and limit the usefulness of such a release device. Thus, we explored the outcome of encapsulation during gelation on the physical state of BSA using CD and SEC. We examined TEA solutions of pH 7.4 and 8.2 as aggregation of BSA is usually attributed to thiol-disulfide interchange in a basic environment (Zhu and Schwendeman 2000
; Maruyama et al. 2001
). Our CD experiments () indicated that the predominantly α-helical structure of BSA (Maruyama et al. 2001
; Abeywickrama et al. 2007
; Watkins et al. 2008
) was preserved independently of pH during gelation as well as after release from the gels.
We also determined via SEC () that BSA did not aggregate upon release from the gel but was instead enriched in monomer solution which was also observed for hyaluronic acid-polyethylene glycol gels (Leach and Schmidt 2005
). Furthermore, there was no significant difference between the % monomer in BSA released at 6 h and at complete gel degradation which suggested that the enriched monomeric content in the released solution was not due to larger aggregates being retained inside the gel. Small fragments were not detected which indicated that hydrolyzed BSA was not released. Polymer-BSA adducts were also not detected in the release solution which confirmed that BSA did not react with PEG during the cross-linking reaction as we and others have shown previously (Elbert et al. 2001
; Zustiak et al. 2010
). Lastly, we concluded that neither TEA nor pH had an effect on BSA aggregation under the examined conditions. This was an important result as gelation time can be controlled effectively by the pH of the gelation solution and a change in pH from 7 to 8 lead to a ~ 3-fold increase in gelation time (Elbert et al. 2001
; Zustiak and Leach 2010
). Control of gelation time has implications in control of the protein environment and minimization of exposure to conditions which over a long time period have the potential to lead to unwanted side reactions between the polymer and protein of interest.