Model for the Analysis of Membrane-Type Dissolution Tests for Inhaled Drugs.

Impactor-type dose deposition is a common prerequisite for dissolution testing of inhaled medicines, and drug release typically takes place through a membrane. The purpose of this work is to develop a mechanistic model for such combined dissolution and release processes, focusing on a drug that initially is present in solid form. Our starting points are the Noyes-Whitney (or Nernst-Brunner) equation and Fick's law. A detailed mechanistic analysis of the drug release process is provided, and approximate closed-form expressions for the amount of the drug that remains in solid form and the amount of the drug that has been released are derived. Comparisons with numerical data demonstrated the accuracy of the approximate expressions. Comparisons with experimental release data from literature demonstrated that the model can be used to establish rate-controlling release mechanisms. In conclusion, the model constitutes a valuable tool for the analysis of in vitro dissolution data for inhaled drugs.

Introduction

In vitro dissolution testing of drug delivery systems is commonly used during the development and manufacturing of dosage forms. Such dissolution tests aid the development and optimization of novel delivery systems and can be used to determine in vitro–in vivo correlations. Moreover, they constitute important quality control tools, but this is currently only utilized for oral drug delivery systems.

Whereas a large number of standardized tests exist for solid dosage forms, no test for orally inhaled products has reached compendial status. However, the development of dissolution tests for delivery systems intended for pulmonary administration has attracted considerable interest, especially during the last decade.[1,2] Although various design principles have been utilized, adequate dispersion of the drug is a common prerequisite. This is generally accomplished by letting the drug deposit onto a membrane, using an impactor such as the Andersen cascade impactor (ACI) or the next-generation impactor (NGI). The membrane is subsequently transferred to the dissolution equipment, and drug release typically takes place through this membrane.

An early dissolution setup for inhalable powders was based on a flow-through design, similar to the USP Apparatus 4, with a recirculating dissolution medium.[3] A solid drug was deposited onto a filter that was transferred to a dissolution cell through which the dissolution medium was pumped. Hence, drug transport across the filter was primarily mediated by convection. Likewise, the standard USP Apparatus 1 (rotating basket) has been used to study release from solid lipid microparticles.[4] Powder samples were wrapped up in sealed glass fiber filters to prevent the microparticles from escaping into the dissolution medium. In this case, drug release is expected to be mediated by a combination of diffusion and convection across the filter.

Many dissolution setups utilize horizontal diffusion cells, as pioneered by Cook et al.[5] Examples include the modified Franz diffusion cell[6,7] and diffusion cells based on permeable Transwell supports.[8,9] Similar designs are utilized for the DissolvIt system[10] and the method proposed by Eedara et al.,[11] but these setups also include a mucus simulant (and the DissolvIt system is reversed since drug release occurs across a membrane placed on top of the drug). In any case, release across the membrane is expected to be primarily mediated by diffusion.

Also, the standard USP Apparatus 2 (paddle) has been adopted for dissolution studies of inhalable powders.[12] In this case, drug particles were deposited onto membranes that were mounted in a cassette that was placed in the USP Apparatus 2. Again, drug release is expected to be primarily mediated by diffusion. Recently, a similar setup has been suggested for use in a Pion μDISS Profiler.[13]

May et al.[14] have proposed a model for drug dissolution, based on the Noyes–Whitney/Nernst–Brunner equation,[15−17] in which consideration was given to the polydispersity of the powders. However, the possible effects of the membrane that separates the donor from the acceptor compartments were not taken into account.

The purpose of this work is to devise a straightforward mechanistic model that combines drug dissolution, described by the mentioned Noyes–Whitney/Nernst–Brunner equation,[15−17] and release across a membrane. Hence, the model is formulated for inhalation powders in which the drug initially is present in solid form. A combined dissolution/release model of this type can be used to determine rate-controlling processes and to analyze experimental release data for which the effect of the membrane cannot be disregarded. Although our basic premise is that release is governed by diffusion, as described by Fick’s first law,[18] the proposed model applies also when convection contributes to the release, provided that drug release is proportional to the concentration of the dissolved drug.

Materials and Methods

Model Formulation

Drug dissolution is assumed to take place in a relatively small donor compartment of volume Vliq that is separated from a considerably larger acceptor compartment by a membrane (Figure ), both containing the same type of dissolution medium. This is considered to be a reasonable assumption, although, in reality, some exchange of the dissolution medium may take place between the compartments. The amounts of the solid and dissolved drug in the donor compartment, expressed as mass or moles per unit volume, are specified in terms of variables S(t) and C(t), where t is the time. All drug is assumed to be present in solid form initially. The dissolved drug is assumed to diffuse through the membrane of surface area Amem and thickness hmem. The effective diffusion coefficient of the drug across the membrane is denoted Dmem. Assuming that the acceptor compartment acts as a sink, conservation of mass requires thatwhere the left-hand side represents the rate of change in the total amount of the drug that remains in the donor compartment and the right-hand side is the (negative) amount of the drug that diffuses across the membrane per unit time, as obtained from Fick’s first law.[18] Notice that both C and S depend on time, although the arguments for notational simplicity are not indicated.

Figure 1

Schematic illustration of the dissolution setup, showing the donor compartment where drug dissolution occurs and the acceptor compartment into which the drug is released by diffusion through the membrane that separates the two compartments.

With the modifications proposed by Nernst[16] and Brunner,[17] the Noyes–Whitney equation[15] can be expressed as[19]where A(t) is the total area of the solid drug, Cs is the solubility of the drug in the dissolution medium, and Dstag is the diffusion coefficient of the drug across a stagnant layer of thickness hstag. Assuming fairly monodisperse particles that retain their shape when undergoing dissolution, the relationship between particle surface area and amount of the solid drug can be expressed as[20,21]where A0 and S0 denote the initial values of A and S, respectively. For simplicity, the thickness of the stagnant layer is here assumed to remain the same throughout the dissolution process, despite the fact that the particles are small.[22] This assumption is considered satisfactory when the particles are located at the membrane boundary. The amount of the drug that has been released through the membrane up to a certain time t is obtained by integration of the drug release rate, i.e., the magnitude of the right-hand-side of eq . Dividing the result by the total amount of the drug initially present in the system, viz., M0 = VliqS0, the fraction of the released drug u(t) is obtained aswhere x is a dummy variable.

Characteristic Time Scale and Nondimensional Form

A characteristic time for diffusion, denoted tdiff, can be defined as followsTo see its physical significance, we assume that all drug has dissolved, so that S = 0. Integration of eq then demonstrates that the concentration of the dissolved drug in the donor compartment decays as e–. In particular, when sink conditions prevail and dissolution is considered to be infinitely fast, the amount of the drug that remains in the donor compartment equals M0 e–, so that the initial release rate is M0/tdiff.

Similarly, a characteristic time for dissolution, denoted tdiss, can be defined as followsTo see its physical significance, we note that eq implies that the initial dissolution rate equals DstagA0Cs/hstag = M0/tdiss since C = 0 initially, in complete analogy with the result obtained above for tdiff.

The ratio between these two time scales will be denoted λ,i.e.,hence, λ ≫ 1 would correspond to a dissolution-limited release and λ ≪ 1 to a diffusion-limited release, i.e., dissolution is the rate-limiting process for λ ≫ 1, whereas diffusion is rate-limiting for λ ≪ 1.

It will be convenient to change the independent variable to τ = t/tdiss. For convenience, we will also introduce the nondimensional variables c = C/S0 and s = S/S0 together with the nondimensional solubility cs = Cs/S0 (remember that S0 is the initial value of S). Hence, cs > 1 if all drug can be dissolved in Vliq and <1 otherwise. Since it is assumed that all drug exists in solid form initially, s = 1 and c = 0 for τ = 0. When expressed in nondimensional variables, eq takes the formCombination of eqs and 3 results inThe fraction of the released drug becomeswhere x is a dummy variable.

Exact Analysis

The balance equation (eq ) is linear in the dependent variables and can therefore be readily integrated with respect to τ. One obtains an expression in terms of the fraction of the released drug, cf. eq , viz.,moreover, eq is separable and can be integrated to producewhere eq has been used. Combination of eqs and 12 results in a nonlinear first-order ordinary differential equation in terms of s(τ)Unfortunately, this equation cannot be solved in closed form. However, as discussed in the following, it forms the basis for an approximate solution procedure. Provided that s(τ) is known, the fraction of the released drug u(τ) can be obtained from eq using the method of integrating factor, the result beingwhereand x is a dummy variable. The integral in eq changes with time τ only when the solid drug remains in the donor compartment, i.e., up to a certain time τ1. This implies that eλτH(τ) and hence 1 + λ eλτH(τ) are constant for all τ > τ1. From eq , the constant value attained by this quantity is obtained as (1 – u1)eλτ, where u1 = u(τ1). Using this result in eq , we obtainfor τ > τ1. Alternatively, eq with s(τ) = 0 may be integrated, subject to the initial condition u(τ1) = u1, to produce the same result. In particular, if cs > 1 and dissolution would proceed infinitely fast so that τ1 = u1 = 0, the fraction of the released drug would be obtained as u(τ) = 1 – e–λτ. A comparison with eq thus reveals that the function H(τ) accounts for the retardation of the release caused by a limited dissolution rate or solubility.

If the fraction of the released drug is to be determined numerically, it is more convenient to combine eqs and 12 to obtain the following nonlinear differential equationThis equation applies as long as the solid drug remains in the donor compartment, i.e., as long as the quantity within the square brackets is positive. It corresponds to eq in the work by Frenning et al.,[23] where a similar procedure was used.

Asymptotic Behaviors

It is instructive to study the behavior of s(τ) when λ ≫ 1 and λ ≪ 1, i.e., the asymptotic behaviors of s(τ) when λ → ∞ and λ → 0.

Dissolution-Controlled Release

When the characteristic time for dissolution is significantly larger than that for diffusion (i.e., when λ ≫ 1), accumulation of the dissolved drug in the donor compartment can be neglected. In this case, eq reduces to a separable ordinary differential equation in s(τ), with solutionwhich is the well-known Hixson–Crowell cube-root law.[20] Clearly, this result is immediately obtained from eq in the limit λ → ∞. Hence, the fraction of the drug that remains in solid form can be expressed asNotice that drug dissolution according to eqs or 19 depends solely on the characteristic time for dissolution since τ = t/tdiss, which is expected when λ ≫ 1.

Diffusion-Controlled Release

When the characteristic time for dissolution is significantly smaller than that for diffusion across the membrane (i.e., when λ ≪ 1), drug release will be biphasic. First, a rapid initial drug dissolution will occur, until either a saturated solution is obtained in the donor compartment or all drug has dissolved, depending on the initial amount of the solid drug present. When the initial drug loading exceeds the solubility (i.e., when cs < 1), a steady state with c = cs will thereafter be maintained as long as the solid drug remains in the donor compartment. Putting dc/dτ = 0 and c = cs in eq , one finds thatimplying thatwhere (1 – cs) represents the drug that remains in solid form after the initial (essentially instantaneous) drug dissolution. The above expression is valid as long as the solid drug remains. Thereafter, eq applies. Notice that the fraction of the drug that remains in solid form according to eq depends on time through the product λτ = (tdiss/tdiff)(t/tdiss) = t/tdiff, i.e., only on the characteristic time for diffusion, as expected when λ ≪ 1.

Sink Conditions

Sink conditions prevail whenever the solubility significantly exceeds the initial drug loading, i.e., when cs ≫ 1. In this case, eqs and 19 continue to be valid, irrespective of the value of λ.

Approximate Analysis

To obtain an approximate solution of the drug release problem, we first seek an expression for the rapid initial drug dissolution obtained when λ ≪ 1 and cs < 1. To this end, we change variable to w = 1 – s in eq . Since w ≪ 1 initially, we keep only the linear terms to obtainwhere μ = 2/3 + λ + 1/cs and ν = 2λ/3. The solution of eq may be expressed in terms of Airy functions[24] but is cumbersome in the subsequent developments. Consistent with the fact that λ ≪ 1 and that we are interested in the initial drug dissolution, we therefore neglect ντ in comparison with μ. For the same reason, we neglect λτ in comparison with 1 on the right-hand side. With these simplifications, the solution of eq subject to the initial condition w(0) = 0 takes the formConsistent with the limiting results expressed by eqs and 21 and the expression for the initial drug dissolution embodied in eq , we postulate the following approximate form for sProvided that τ1 is known, the constants A, B, and a can be determined fairly conveniently from the conditions that the functional value and the first two derivatives of s̃ attain the correct values at τ = 0. The derivations are provided in Appendix A. It turns out to be more involved to determine an accurate value for τ1 itself, which represents the value of τ at which dissolution is complete. Our approach is based on demanding that eq be satisfied when τ = τ1. This condition can be translated to a nonlinear algebraic equation from which τ1 can be determined. Again, the derivations are summarized in Appendix A.

When evaluating the integral that results when s̃ is substituted for s in eq , we proceed slightly differently depending on the value of the constant a. When a ≠ 0, we change variable to σ = 1 – aτ/τ1 and let b = 1 – a, so thatwhere x is a dummy variable. For convenience, we have used the shorthand notation λ1 = λτ1/a and μ1 = μτ1/a. It proves convenient to introduce the auxiliary function

where the constant of integration has been omitted and where f(x) is a function defined aswith Ei being an exponential integral.[24] Moreover, to evaluate the integral in eq , partial fraction decomposition of the rational function in x and integration by parts have been performed. Using eq in eq , we obtainwhere K1 = AF(λ1 – μ1, b; 1) + BF(λ1, b; 1) is an auxiliary constant that represents the contribution from the lower endpoint of the integral in eq . Visual basic routines that can be used to determine the functions f(x), F(α, b; x) and H(τ) using, e.g., Microsoft Excel are provided in the Supporting Information.

The above analysis cannot be used when a = 0 (or very small) since a appears in denominators. Such a values are to be expected when sink conditions prevail since s(τ) then approaches the asymptotic form given by eq , which is identical in form to s̃ obtained from eq when a = A = 0. In the special case of a = 0 (or very small), we instead change variable to ρ = 1 – τ/τ1 to obtainwhere λ2 = λτ1 and μ2 = μτ1 and x again is a dummy variable. Defining the auxiliary functionwhere the constant of integration has been omitted, we may in analogy with eq writewhere K2 = AG(λ2 – μ2; 1) + BG(λ2; 1) is an auxiliary constant that represents the contribution from the lower endpoint of the integral in eq . The fraction of the released drug is finally obtained using H(τ) as obtained from eq or 31 in eq . Visual basic routines that can be used to determine the functions G(α; x) and H(τ) using Excel are provided in the Supporting Information.

Numerical Evaluation

The standard Newton–Raphson method, with an initial value of 1/3, was used to solve eq for ω = 1/τ1 (cf. Appendix A and Supporting Information). The function H(τ) was evaluated by eq when a > 0.01 and by eq otherwise. The exponential integral was evaluated using an Excel VBA implementation of the algorithm described by Paciorek (see the Supporting Information).[25] The approximate analytical solutions were compared to numerical solutions obtained using the Runge–Kutta Fehlberg method implemented in Maple 2019.1 (Maplesoft, Canada).

Results and Discussion

Assessment of Accuracy and Parametric Study

Model calculations were performed for three different values of the nondimensional solubility cs (0.1, 1, and 10), defined as the ratio between the solubility of the drug in the donor compartment and the initial drug loading. For each value of cs, three different values of the ratio λ (0.1, 1, and 10) between the characteristic times for dissolution and diffusion were considered. The results obtained are summarized in Figure . Specifically, Figure a–c displays the fraction of solid drug s remaining in the donor compartment, and Figure d–f shows the fraction of drug u that has been released to the acceptor compartment. The results obtained for cs = 0.1 are shown in Figure a,d, for cs = 1 in Figure b,e, and for cs = 10 in Figure c,f.

Figure 2

Fraction of solid drug s remaining in the donor compartment (a–c) and the fraction of the released drug u (d–f) vs nondimensional time τ as obtained analytically (solid lines) and numerically (dashed lines). Model calculations were performed for cs = 0.1 (a, d), cs = 1 (b, e), and cs = 10 (c, f). In each case, results are presented for λ = 0.1, 1 and 10. For comparison, the limiting results obtained when λ ≪ 1 and λ ≫ 1 are included in the top-left graph (dotted lines).

Before discussing the individual cases in detail, we note that the analytical approximation (solid lines) generally exhibits a good to excellent correspondence with the numerical results (dashed lines). Minor deviations can be seen especially when diffusion across the membrane has a significant impact on the dissolution and release profiles. This is expected because drug dissolution and release by diffusion become more interrelated in this case. However, also, in this case, an adequate correspondence is observed. The analytical approximation has been tested in a wider parameter range, and its correspondence to the analytical solution remains adequate. For example, the maximal absolute errors observed in the parameter range 0.01 ≤ cs, λ ≤ 100 were 0.033 for s and 0.022 for u. Hence, we can conclude that the analytically derived expressions can be used with confidence, e.g., in the analysis of experimental release data (cf. Section below).

Figure a,d corresponds to dissolution and release when no more than 10% of the initial drug loading can be dissolved in the donor compartment. Depending on the value of λ, drug release ranges from being largely dissolution-controlled (for λ = 10) to being largely diffusion-controlled (for λ = 0.1). This is corroborated by a comparison with the limiting results obtained for λ ≪ 1 and λ ≫ 1, embodied in eqs and 21, which are included in Figure a (dotted lines). It can be noted that a value of λ > 10 is needed to obtain a complete dissolution control since the solubility is low. On the contrary, the agreement with the diffusion-controlled result is almost perfect when λ = 0.1. In this case, a rapid initial drug dissolution is clearly seen in the fraction of the solid drug remaining in the donor compartment, cf. eq , which corresponds to a small delay in the fraction of the drug being released. Since only a small fraction of the total amount of the drug can be dissolved in the donor compartment, the fraction of the released drug closely mirrors the amount of the drug being dissolved.

Figure b,e corresponds to dissolution and release when all of the initial drug loading (but no more) can be dissolved in the donor compartment. Drug release is dissolution-controlled for λ = 10, and the limited solubility in the donor compartment clearly affects dissolution for smaller values of λ. However, this effect is not as pronounced as it was for cs = 0.1, as expected.

Figure c,f corresponds to dissolution and release when a dose 10 times larger than the one present can be dissolved in the donor compartment. Since dissolution occurs under sink conditions, the results obtained for the fraction of the solid drug remaining in the donor compartment collapse on a single curve, corresponding to the limiting result expressed by eq . The parameter λ nevertheless has a decisive influence on the fraction of the released drug since it controls the rate at which the dissolved drug diffuses across the membrane.

Comparison with Experimental Release Data

When expressed in the nondimensional form, the fraction of the released drug depends on drug solubility in relation to the initial drug loading (cs = Cs/S0) and the ratio between the time scales for dissolution and diffusion (λ = tdiss/tdiff). For real release data, the time scale for dissolution (tdiss) is needed to convert dimensional time to nondimensional time according to τ = t/tdiss. Hence, the fractional release u(t) generally depends on three parameters, which, e.g., may be selected as tdiss, tdiff, and cs = Cs/S0. The situation is different when sink conditions prevail, however, since these parameters are dependent. The reason for this is that dissolution no longer is impeded by a limited solubility, implying that u(t) becomes independent of cs. The dissolution profile is well described by the asymptotic equation (eq ) whenever cs ≫ 1, which is identical in form to the profile described by eq if a = A = 0 (and consequently B = 1) and τ1 = 3. Hence, eq reduces toso that, according to eq As seen from the above equation, the fraction of the released drug thus depends on time via the two quantities τ = t/tdiss and λτ = t/tdiff. Hence, tdiss and tdiff between them define the release profile and no value for cs needs to be provided.

Moreover, an incomplete release is often seen in experimental release data, i.e., the fraction of the released drug levels out at a value smaller than 100%. This may be caused by a loss of the drug during deposition or binding of the drug to surfaces or materials during the release experiment. To accommodate for this phenomenon in a simplified manner, the theoretical amount of the released drug can be multiplied by a constant N < 1.

Data from Sakagami et al.

Sakagami et al. have recently presented dissolution data for a range of orally inhaled corticosteroid products using a Transwell-based dissolution setup.[26] The dissolution medium consisted of 10 mL of simulated lung-lining fluid with 0.02% w/v dipalmitoyl phosphatidylcholine (DPPC) in the donor compartment. For the more soluble corticosteroids flunisolide (FN), triamcinolone acetonide (TA), and budesonide (BUD), the same dissolution medium was used in the acceptor compartment. However, for the less soluble ones, 1% w/v d-α-tocopheryl poly(ethylene glycol) 1000 succinate (TPGS) was added to the dissolution medium in the acceptor compartment. The addition of TPGS to only one compartment makes the interpretation of the results more difficult since some movement of the dissolution medium between the compartments likely occurs. For this reason, we here focus on the more soluble steroids, for which the dissolution profiles presented in Figure were obtained.[26] The parameter cs was calculated from data obtained from Sakagami et al.[26] and N was set equal to 1 since a complete dissolution was observed for all three drugs (Table ). The characteristic time for diffusion across the membrane (tdiff) was determined from data for FN (dotted curve), for which dissolution was found to be very rapid.[26] Varying the characteristic times for dissolution (tdiss) while keeping tdiff fixed at the value obtained for FN resulted in the dashed and solid lines corresponding to TA and BUD, respectively. For numerical reasons, tdiss was required to be at least 10–3 min (indicated as ≪1 min in Table ). The agreement between the theoretical release profiles and the experimental data is considered satisfactory, especially since notable variations between repeated experiments were observed by Sakagami et al.[26] These results indicate that the difference seen between the drugs indeed can be attributed to differences in their solubility.

Figure 3

Comparison between experimental data (symbols) from Sakagami et al.[26] and model calculations (dotted, dashed, and solid lines) using the parameters collected in Table for drugs flunisolide (FN), triamcinolone acetonide (TA), and budesonide (BUD).

Table 1

Parameters Used in the Model Calculations Presented in Figure for Drugs Flunisolide (FN), Triamcinolone Acetonide (TA), and Budesonide (BUD)

substance	cs (−)	tdiss (min)	tdiff (min)	N (−)
FN	17.9a	≪1	88.0	1
TA	4.64a	75.5	88.0b	1
BUD	4.29a	97.0	88.0b	1

Calculated from data obtained from Sakagami et al.[26]

Kept constant at the value obtained for FN.

Data from Eedara et al.

Eedara et al. have recently described a novel dissolution method for inhaled drugs that consists of a donor compartment that is separated from a flow-through cell by a dialysis membrane.[11] The method was used to investigate dissolution and diffusional transport of the antitubercular drugs moxifloxacin and ethionamide using phosphate-buffered saline (PBS) as the dissolution medium. Poly(ethylene oxide) (PEO) was used to simulate the mucus layer. The results obtained for ethionamide exhibited a release profile that resembled the classical Higuchi square-root-of-time law,[27,28] indicating that other mechanisms than those considered in this work may be at work. Examples of release profiles obtained for moxifloxacin at different perfusion rates, as obtained by Eedara et al.,[11] are presented in Figure . Some data at larger times have been excluded to more clearly display the initial delay. Moreover, the release profiles exhibited a gradual convergence toward complete release for larger times (data not shown), which here is interpreted as resulting from a slow release of the drug embedded in the PEO matrix. For this reason, we focus on the initial release and put N equal to 0.9. The parameter cs (Table ) was calculated from data provided by Eedara et al.[11] The release profiles in Figure exhibit a dependence on the permeation rate and seem to converge for sufficiently high rates. As noted by the authors, this behavior is consistent with the one expected from the presence of an unstirred water layer, which effectively acts as external mass-transfer resistance.[29,30] For this reason, the characteristic time for dissolution (tdiss) was determined from data for a perfusion rate of 0.8 mL/min (solid curve in Figure ). Varying the characteristic times for diffusion (tdiff) while keeping tdiss fixed at the value obtained for 0.8 mL/min resulted in the dotted and dashed lines corresponding to the perfusion rates of 0.2 and 0.4 mL/min, respectively. Again, the correspondence between theory and experiments is considered satisfactory, especially since there are inherent uncertainties in the experimental data, corroborating the interpretation that the differences seen between perfusion rates are due to an external mass-transfer resistance (unstirred water layer).

Figure 4

Comparison between experimental data (symbols) from Eedara et al.[11] and model calculations (lines) using the parameters collected in Table .

Table 2

Parameters Used in the Model Calculations Presented in Figure

permeation rate (mL/min)	cs (−)	tdiss (min)	tdiff (min)	N (−)
0.2	8.84a	4.52b	22.9	0.9
0.4	8.84a	4.52b	13.5	0.9
0.8	8.84a	4.52	13.1	0.9

Calculated from data obtained from Eedara et al.[11]

Kept constant at the value obtained for a permeation rate of 0.8 mL/min.

Data from Rohrschneider et al.

Rohrschneider et al. have presented dissolution and release data for the three inhaled corticosteroids BUD, ciclesonide (CIC), and fluticasone propionate (FP) using a Transwell-based dissolution setup.[31] To reduce the diffusional resistance, the polycarbonate membrane was removed and the glass microfiber filter/filter paper onto which the drug was deposited was placed on thermoformed notches in the Transwell insert. Dissolution experiments were generally performed using 0.5% w/v sodium dodecyl sulfate (SDS) in the PBS solution. Examples of the results obtained are shown in Figure .[31] The parameter cs was calculated from data obtained from Rohrschneider et al.[31] The parameter N was set to 1 for BUD and CIC since a nearly complete release occurred for these drugs. The remaining parameters were determined by curve fitting, see Table . Hence, the characteristic times for dissolution and diffusion (tdiss and tdiff) were both allowed to vary. However, for numerical reasons, both characteristic times were required to be at least 10–3 min (indicated as ≪1 min in Table ). The results are shown by solid lines in Figure .

Figure 5

Comparison between experimental data (symbols) from Rohrschneider et al.[31] and model calculations (solid lines) using the parameters collected in Table for drugs budesonide (BUD), ciclesonide (CIC), and fluticasone propionate (FP). Fraction of the released drug (left) and a semilog plot of the fraction of the remaining drug (right).

Table 3

Parameters Used in the Model Calculations Presented in Figure for Drugs Budesonide (BUD), Ciclesonide (CIC), and Fluticasone Propionate (FP)

substance	cs (−)	tdiss (min)	tdiff (min)	N (−)
BUD	2.35a	1.26	33.9	1
CIC	2.00a	≪1	66.8	1
FP	0.20a	158	≪1	0.65

Calculated from data obtained from Rohrschneider et al.[31]

The model results suggest that release is diffusion-controlled for both BUD and CIC for which cs > 1 so that all drug can be dissolved in the donor compartment. This result may appear somewhat surprising since a clear difference between release of the drug and release of the solution was observed by Rohrschneider et al.[31] but can be explained by differences in λ, i.e., the ratio between the characteristic time scales for dissolution and diffusion. That diffusion dominates can also be seen in Figure b, which displays the fraction of the remaining drug (i.e., 1 – u) on a logarithmic scale. If all drug dissolves rapidly in the donor compartment, the fraction of the remaining drug is expected to decrease e–/ [see the discussion following eqs and 16 with u1 = τ1 = 0], so that a linear decrease would be seen in a semilog plot. This is indeed found, especially for CIC but also for BUD (notice that deviations from a linear relationship may occur at low fractions of the remaining drug because less than 100% of the drug may be released). It would be tempting to attribute the differences seen in the release rate between BUD and CIC to differences in solubility. However, provided that cs > 1 and dissolution is rapid, other factors are likely involved since the initial release rate then equals M0/tdiff, where M0 is the total amount of the drug present in the system and tdiff only depends on the diffusion coefficient across the membrane and geometrical factors [eq and the discussion following that equation]. Although a higher solubility does result in a higher release rate in terms of grams or moles per second, a larger amount of the drug needs to be transported, so that the fractional release rate will be independent of solubility.

The situation is different for FP, where our data suggests that the release is completely dissolution-controlled and that only about 65% of the drug is released. An incomplete release of FP is consistent with other data presented by Rohrschneider et al.[31] and may, for example, be the result of losses of the drug due to its binding to surfaces in the dissolution setup.

Conclusions

A model of dissolution based on the well-established Noyes–Whitney/Nernst–Brunner equation coupled with diffusional transport across a membrane was formulated and expressed in nondimensional form. A closed-form analytical approximation was derived. This approximation has sufficient accuracy to be used with confidence irrespective of the rate-controlling mechanism(s). The usefulness of the model to establish rate-controlling mechanisms was demonstrated by comparisons with experimental release data obtained from the literature.