Preclinical Comparison of an Injectable In-Situ Scaffold-Forming Oligomeric Collagen for Soft Tissue Volume Augmentation in a Rat Model

Abstract

Introduction:

Dermal fillers are widely used to restore facial volume, yet current materials often exhibit swelling, inflammatory responses, or limited persistence.

Objective:

To compare rheological and extrusion properties, volumetric persistence, and local tissue response of an injectable, in-situ scaffold-forming oligomeric collagen with a commercially available hyaluronic acid filler and saline in a rat model.

Methods:

Oligomeric collagen was administered by subcutaneous injection, with commercial filler and saline as controls. Animals were evaluated at 1, 6, and 12 weeks using three-dimensional volumetric scanning and histopathological analyses. Rheological, compressive, and extrusion properties were measured. Statistical comparisons used one-way analysis of variance or Student’s t-tests.

Results:

In this 12-week rat study, oligomer demonstrated low extrusion force and stable scaffold volumes after equilibration. Extrusion force was similar to saline and ∼15-fold lower than the commercial filler. Scaffold volumes were ∼25% of the injected volume and remained stable from 1 to 12 weeks. The commercial filler exhibited swelling (∼400%) at 1 week, followed by ∼83% volume loss. Histological analysis showed minimal inflammatory response with scaffold integration, whereas the commercial filler exhibited regional inflammation and fibrous capsule formation.

Conclusions:

Injectable scaffold-forming oligomer exhibited low extrusion force, persistent volume, and minimal local tissue response over 12 weeks.

Key Points

Question: How does a new collagen filler that forms a scaffold after injection compare with a hyaluronic acid (HA) filler and saline in terms of injectability, volume stability, and tissue response in an animal model?

Finding: Compared with the HA filler, the collagen filler required less force to inject, maintained stable volume over 12 weeks, remained localized without capsule formation, and showed minimal inflammation similar to saline.

Meaning: This approach may offer a more predictable and longer-lasting option for soft tissue volume restoration.

Introduction

Skin aging is a multifaceted process involving intrinsic and extrinsic mechanisms that lead to progressive structural changes in skin and underlying tissues.¹ Loss and disorganization of collagen contribute to wrinkle formation, reduced skin firmness, and volume loss.² Dermal fillers are widely used to restore or modify soft tissue volume and facial contours, thereby rejuvenating aging skin or correcting soft tissue defects. Dermal filler injections are among the most frequently performed nonsurgical aesthetic procedures worldwide, with over 9.3 million procedures reported in the United States in 2024.³

Currently available dermal fillers span a range of material classes, including absorbable (temporary) fillers such as collagen, hyaluronic acid (HA), calcium hydroxyapatite (CaHA), and poly-l-lactic acid (PLLA), as well as nonabsorbable (permanent) fillers such as polymethylmethacrylate (PMMA). Despite widespread use, these materials remain limited by challenges in injectability, volumetric predictability, durability, and tissue integration.^4–8 Hydrophilic expansion can lead to early swelling, while material migration may compromise spatial control and contour precision.^8–10 Crosslinked or particulate formulations may provoke foreign body inflammatory responses and fibrous encapsulation, and progressive bioresorption can result in gradual loss of correction and the need for repeat procedures.^10–12 Fillers are also associated with pain, erythema, and swelling due to tissue trauma, inflammation, and fluid absorption.^6,7,13,14 Additional complications include surface irregularities, granuloma formation, infection, allergic reactions, local tissue necrosis, and material migration. Although rare, intravascular injection can lead to serious events such as skin necrosis, vision loss, or stroke.^15–17 Collectively, these limitations highlight the need for next-generation fillers that improve procedural control, volumetric stability, and tissue response.

Given the central role of fibrillar collagen in tissue structure and mechanics, collagen was the first material widely adopted for dermal filler applications.¹⁸ As summarized in Supplementary Table S1, collagen-based fillers are primarily composed of particulate fibrillar collagen derived from animal tissue or cultured fibroblasts.^19–23 Chemical crosslinking (e.g., glutaraldehyde, d-ribose) has been employed to improve firmness and longevity; however, exogenous crosslinkers are associated with increased foreign body responses and inflammation.^24–26 Bovine-derived collagen fillers historically also required pretreatment skin testing due to hypersensitivity risk.²⁷ Despite limitations related to persistence and immunogenicity, collagen fillers dominated the market for more than two decades.

HA-based fillers are now the most widely used in North America.²⁸ HA is a naturally occurring polysaccharide sourced from bacterial fermentation or animal tissues and is formulated at different molecular weights and chemically crosslinked to yield products with varying physicochemical properties, longevity, and intended use.²⁹ The volumizing effect of HA fillers typically persists for 6–24 months, depending on the formulation and application site.^30,31 Biostimulatory fillers composed of microparticles (e.g., CaHA, PLLA, PMMA) represent an alternative approach, augmenting soft tissue volume through induction of host collagen and extracellular matrix deposition. These materials induce varying degrees of foreign body inflammatory responses, including fibrous encapsulation and collagen deposition.^32–34

Our approach to next-generation soft tissue fillers leverages a polymerizable type I oligomeric collagen (oligomer) that forms fibrillar scaffolds in situ. This highly purified type I collagen building block is extracted and purified from porcine dermis under conditions that remove cellular and immunogenic components.³⁵ When acidic oligomer solutions are mixed with a self-assembly reagent (buffer) and brought to physiological pH and ionic strength, polymerization is initiated, resulting in rapid formation of a stable, viscoelastic fibrillar scaffold without exogenous crosslinking agents (Fig. 1).^35,36 Prior gross, histological, and transcriptomic analyses across multiple tissue environments demonstrate that oligomer-based materials exhibit immunotolerance, avoiding foreign body responses, fibrous capsule formation, and immune-mediated bioresorption (biodegradation).^37–47 These scaffolds integrate with surrounding tissues, provide durable mechanical support, and undergo regenerative remodeling characterized by physiological collagen turnover with mechanobiologically driven formation of site-appropriate functional tissue.^46,47 By enabling low-viscosity injection followed by rapid in-situ fibrillar assembly without exogenous crosslinkers, this platform is designed to enhance procedural control, reduce swelling and migration, support tissue integration, and mitigate inflammatory-mediated encapsulation observed with existing technologies.

Fig. 1.

Overview of injectable, in-situ scaffold-forming oligomer. (A) Schematic illustrating oligomer phases before, during, and after polymerization. Prior to polymerization, oligomer is a liquid comprising acid-soluble type I oligomeric collagen molecules. Mixing with the self-assembly reagent (buffer) adjusts the solution to physiological pH and ionic strength, initiating polymerization characterized by nucleation and early fibril self-assembly. Following polymerization (approximately 30–60 s), oligomer forms a viscoelastic scaffold composed of an insoluble fibrillar network with interstitial fluid. (B) Photograph showing mixing of acidic oligomer and buffer solutions via a Luer-lock syringe connector. (C) Scanning electron micrograph of in-situ polymerized fibrillar scaffold following in vivo subcutaneous injection. (D) Ex vivo images of mixed oligomer solution after injection into a plastic mold, demonstrating the transition from a translucent solution to an opaque, shape-retaining fibrillar scaffold.

In this study, we evaluated an injectable scaffold-forming oligomeric collagen in comparison with a commercially available HA filler and saline. We hypothesized that in-situ scaffold formation would enable low-force extrusion, stable volumetric behavior, and a minimal local tissue response. Material properties were characterized using oscillatory shear rheometry, unconfined compression, and extrusion force measurements, and in vivo performance was assessed through three-dimensional (3D) volumetric scanning and histological evaluation of skin explants following subcutaneous injection at 1, 6, and 12 weeks.

Materials and Methods

Injectables

Polymerizable type I oligomeric collagen (oligomer) was obtained as a sterile two-component kit (acidic oligomer solution and self-assembly buffer) from GeniPhys, Inc., Indianapolis, IN. Solutions were aseptically loaded into syringes (BD, Franklin Lakes, NJ). Juvéderm Ultra Plus XC (Allergan, Irvine, CA) and phosphate buffered saline (PBS; ThermoFisher Scientific, Waltham, MA) served as commercial filler and saline controls, respectively.

Rheometric analysis

Viscoelastic and compressive properties were measured using an HR30 rheometer (TA Instruments, New Castle, DE) with a 40-mm parallel plate geometry. Oscillatory shear testing was performed at 1% strain and 1 Hz (725 µm gap; 37°C) to determine storage modulus (G′), shear loss modulus (G″), and tanδ (G″/G′). Compression modulus was calculated from unconfined compression stress-strain curves generated by applying compressive strain at 2.76%/s and determining the slope within the low-strain region (20–40% strain). Time-dependent oscillatory shear measurements were used to assess oligomer polymerization kinetics.⁴²

Extrusion force measurement

Extrusion force testing was performed in accordance with ISO 7886-1⁴⁸ using an Instron Universal Testing System (Model 68SC-5; Norwood, MA) with a custom syringe clamp (Supplementary Fig. S1A). Injectables were loaded into 1-mL syringes (BD) fitted with 26-gauge, 3/8-inch needles (BD). Plunger displacement was applied at 100 mm/min and terminated with 0.1 mL remaining. Maximum (F_max) and average (F_avg) forces were defined as the peak over 0–3 mm and the mean over 5–15 mm of displacement.

Rat subcutaneous injection model

All animal procedures were approved by the Purdue University Institutional Animal Care and Use Committee and conducted in accordance with AAALAC guidelines. Lewis rats (250–275 g; Envigo, Indianapolis, IN) received subcutaneous injections (n = 4 per group per timepoint) at four standardized dorsal sites using a 26-gauge, 3/8-inch needle (BD). Oligomer (1 mL) was mixed immediately prior to injection using a Luer-lock connector (Baxter Healthcare, Deerfield, IL), with saline (1 mL) as a control. A reduced volume (0.5 mL) of the commercial filler was administered to account for hydrophilic swelling¹⁰ and maintain discrete injection sites. Post-injection evaluations were performed at 1, 6, and 12 weeks.

3D scanning

For volumetric analysis, excised skin containing each injection site was scanned using a Creaform Go!SCAN SPARK handheld 3D scanner (Quebec, Canada) with VXElements software (version 10.0.1.10229). Post-processing and quantitative analysis (Supplementary Fig. S2) were performed using VXModel software (version 10.0.1.10229).

Gross and histological analyses

Tissue explants were paraffin-embedded and stained with hematoxylin and eosin (H&E). Selected sections were immunostained with mouse anti-rat CD68 (pan macrophage marker; 1:100; Bio-Rad, Hercules, CA) or rabbit anti-CD11b (pan leukocyte marker; 1:60,000; Abcam, Waltham, MA), followed by detection using ImmPRESS HRP Anti-Mouse or Anti-Rabbit IgG Polymer Kits (Vector Laboratories, Burlingame, CA) and hematoxylin counterstaining. Histological assessments were performed by an independent pathologist blinded to treatment groups. Whole-slide images were acquired using an Aperio VERSA 8 scanner (Leica Biosystems, Deer Park, IL).

Statistical analysis

Data were analyzed using Prism 6 (GraphPad Software, Boston, MA). Multi-group comparisons were performed using one-way analysis of variance (ANOVA) with Tukey HSD post-hoc testing, and two-group comparisons were performed using two-tailed Student’s t-tests. Statistical significance was defined as p < 0.05.

Results

Extrusion force, rheological, and compressive properties of injectables

Oligomer and saline were low-viscosity solutions with minimal force variability compared with the commercial filler. Oligomer F_max and F_avg values were 1.72 ± 0.15 N and 1.65 ± 0.21 N, respectively, and were not statistically different from saline (F_max: 0.80 ± 0.06 N; F_avg: 0.66 ± 0.11 N), while being approximately 15-fold lower than those of the commercial filler (Table 1, Supplementary Fig. S1). Once mixed, oligomer polymerized in 38.8 ± 3.3 s at 37°C. The resulting oligomer scaffold exhibited rheological and compressive properties (Table 1), with G′ and E_comp values of 3830 ± 134 Pa and 9.03 ± 0.90 kPa, respectively. In contrast, the commercial filler, which remained a gel during and after injection, exhibited lower stiffness, with G′ values of 148 ± 13 Pa. Tanδ values were similar for oligomer and commercial injectables, and E_comp values for the commercial filler were modestly higher.

Table 1.

Summary of extrusion force, rheological, and compressive properties for oligomer and commercial injectables

Injectable	F_max; F_avg (N)	G′ (Pa)	G″ (Pa)	Tanδ = G″/G′	E_comp (kPa)
Commercial	29.11 ± 1.29^b; 23.57 ± 0.78^b	148 ± 13^a	27.5 ± 2.5^a	0.188 ± 0.010^a	11.2 ± 1.1^a
Oligomer	1.72 ± 0.15^a; 1.65 ± 0.21^a	3830 ± 134^b	524 ± 21^b	0.137 ± 0.006^b	9.03 ± 0.90^b
Saline^*	0.80 ± 0.06^a; 0.66 ± 0.11^a	N/A	N/A	N/A	N/A

Values sharing the same superscript letter within a column are not significantly different (one-way ANOVA with Tukey post-hoc test for multiple group comparisons; unpaired two-tailed t-test for two-group comparisons; p < 0.05).

Saline was tested under identical extrusion conditions. Rheological (G′, G″, Tanδ) and compressive modulus (E_comp) values are not applicable under unconfined compression due to its Newtonian fluid behavior.

ANOVA, analysis of variance.

Subcutaneous injection

All injectables were well tolerated, with no evidence of erythema or other adverse skin changes. Immediately following administration, all injectables formed discrete, palpable regions. Within 24 h, oligomer sites showed localized fullness that was slightly reduced in size. In contrast, regions injected with the commercial filler increased in volume over the first several days, consistent with swelling. No visible bulking was noted for saline-injected sites after the first 24 h. Bulked regions associated with both oligomer and commercial injectables remained identifiable throughout the study period by visual inspection and palpation, with commercial sites appearing more prominent. The commercial filler exhibited displacement from injection sites, resulting in merged rostral regions evident at 1 week (Fig. 2A, Commercial, 1 week) that persisted at later time points.

Fig. 2.

Qualitative and quantitative volumetric analyses of subcutaneous injection sites showing stable oligomer geometry between 1 and 12 weeks. (A) Representative photographs of injection sites following saline, commercial filler, and oligomer injection at 1-, 6-, and 12-week time points. The commercial filler image at 1 week shows the merging of rostral sites, as observed in all animals within this group. Quantitative 3D scanning analysis of explanted injection sites showing (B) volume, (C) perimeter, (D) surface area, and (E) height for commercial filler and oligomer groups. For each group, letters denote statistically significant differences (p < 0.05). Sample sizes were n = 4 for oligomer and n = 2 for commercial filler, as merged material sites were excluded from quantitative analysis.

Gross examination and volumetric analysis of 3D-scanned explanted injection sites (Fig. 2) revealed readily identifiable scaffolds at all time points for the oligomer group. These scaffolds appeared whitish and showed gross evidence of integration with the surrounding connective tissue (Fig. 2A). At 1 week, measured scaffold volumes were lower (∼25%) than the injected volume. Quantitative analysis revealed no significant changes in scaffold volume, perimeter, surface area, or height between 1 and 12 weeks (Fig. 2B–E), indicating stabilization of scaffold geometry after the initial equilibration period.

In contrast, the commercial filler was identified as a yellowish mass surrounded by a fibrous capsule and exhibited minimal tissue integration. Disruption of the capsule resulted in leakage of the viscous material from the injection site. Quantitative assessment demonstrated initial swelling at 1 week, followed by significant time-dependent decreases in volume (82%), perimeter (51%), and surface area (70%). Changes in height over time were modest but not statistically significant (Fig. 2E). No visual or quantifiable material was detected at saline-injected sites at any time point (Fig. 2).

Histological outcomes

Histological analysis demonstrated that scaffolds formed by the oligomer injectable were readily identifiable as eosinophilic material located immediately beneath the panniculus carnosus muscle (Fig. 3C–D). Oligomer and saline injection sites exhibited similar tissue responses characterized by minimal inflammatory cell infiltration (Fig. 3). Immunohistochemical staining for CD11b-positive leukocytes and CD68-positive macrophages demonstrated similar qualitative staining patterns between oligomer and saline groups (Fig. 4).

Fig. 3.

Histological assessment of subcutaneous injection sites showing persistence of the oligomer scaffold with tissue responses consistent with immunotolerance. Representative H&E-stained cross-sections of (A) saline, (B) commercial filler, and (C) oligomer injection sites at 1-, 6-, and 12-week time points. Scale bars: 1 mm (left panels) and 200 µm (right panels). CF, commercial filler; D, dermis; FC, fibrous capsule; OS, oligomer scaffold; PC, panniculus carnosus muscle; SQ, subcutaneous tissue.

Fig. 4.

Immunolabeling of inflammatory cell markers showing tissue responses at oligomer injection sites consistent with minimal immune reactivity and comparable to saline. Representative tissue cross-sections of saline, commercial filler, and oligomer injection sites at 1-, 6-, and 12-week time points stained for (A) CD11b (pan-leukocyte marker) and (B) CD68 (pan-macrophage marker). Scale bars: 100 µm.

Integration of the oligomer scaffold with the surrounding tissue was observed as stromal cell infiltration at the scaffold-tissue interface at 1 week (Supplementary Fig. S3A). By 12 weeks, vascular structures were present within the scaffold, along with increased cellularization, and small regions of adipose tissue were observed within a subset of scaffolds, while substantial regions of nonremodeled oligomer scaffold remained (Supplementary Fig. S3B). The commercial filler exhibited regional inflammatory cell infiltration, particularly at the material periphery, along with fibrous capsule formation that was evident as early as 1 week (Fig. 3). Time-dependent reductions in material volume were evident from tissue cross-sections between 1 and 12 weeks.

Discussion

The present findings demonstrate that an in-situ scaffold-forming oligomeric collagen exhibits handling, volumetric, and tissue response characteristics distinct from a commercial HA filler. Following subcutaneous injection in a rat model, the material transitioned from a low-viscosity solution to a stable fibrillar scaffold that remained localized and supported tissue integration without fibrous encapsulation. These observations are consistent with our hypothesis that in-situ scaffold formation enables controlled delivery, stable volumetric behavior following initial equilibration, and a tissue-integrating response in this model.

The injectability of dermal fillers is an important parameter influencing procedural ease, dosing precision, and tissue trauma.^13,49 The phase-changing nature of the oligomer injectable supported consistent extrusion with minimal force variability, in contrast to what has been reported for particulate filler formulations.⁴ Extrusion forces for oligomer solution following mixing were similar to saline and approximately 15-fold lower than those of the commercial filler and within ISO 7886-1 thresholds (≤2 mL syringes; maximum allowable forces of 10 N [F_max] and 5 N [F_avg]).⁴⁸ The material transitioned from a low-viscosity solution to a stable fibrillar scaffold within 30–60 s.

After polymerization, the oligomer scaffold exhibited rheological and compressive properties within the range reported for human skin, behaving as an elastic solid rather than a gel. Reported G′ values for human skin range from approximately 1600–4900 Pa.⁵⁰ The measured G′ (3830 ± 134 Pa), G″ (524 ± 21 Pa), and tanδ (0.137) values indicate stability under compression, shear, and torsional forces relevant to soft tissue. In comparison, G′ values for conventional HA fillers typically range from 10 to 1000 Pa, with increasing G′ often associated with increased molecular weight, concentration, and/or crosslinking, increased gel viscosity, and increased extrusion force.⁵¹ Importantly, the oligomer scaffold exhibits viscoelastic properties that fall within the range reported for native skin while maintaining injection forces comparable to saline, a combination not typically observed in current dermal fillers.

In vivo, the oligomer scaffold demonstrated sustained volumization from 1 to 12 weeks and remained localized at the injection site without evidence of migration. Scaffold remodeling, including cellularization and vascularization, was consistent with prior studies in which scaffold-forming oligomer materials were administered in subcutaneous, intramuscular, and endoscopically guided vocal fold applications.^40,41,43,52 The oligomer injectable was not associated with foreign body responses or immune-mediated bioresorption observed with temporary fillers or fibrotic responses reported for some biostimulatory fillers.^4,11,12 Quantitative 3D analysis revealed that the scaffold volume equilibrated to ∼25% of the injected volume by 1 week, consistent with loss of interstitial fluid during scaffold equilibration. No subsequent changes in volume or geometry were detected through 12 weeks, indicating persistence. By contrast, the commercial filler exhibited early swelling (∼400% of the injected volume at 1 week) followed by volume loss over time (∼83% between 1 and 12 weeks), consistent with previous preclinical and clinical observations.^10,21 Given the importance of volumetric predictability,^5,6 these findings highlight the potential advantages of an in-situ scaffold-forming material that minimizes swelling, migration, and rapid biodegradation.

Consistent with prior clinical experience using porcine-derived collagen fillers, which have not required routine pretreatment skin testing,^53,54 no evidence of hypersensitivity was observed in this preclinical model. In the context of dermal fillers, reversibility remains an important design consideration. While HA fillers are routinely reversed using hyaluronidase, collagen-based materials may offer a complementary approach through collagenase-mediated degradation. Supporting this concept, prior work has demonstrated that oligomer scaffolds were susceptible to degradation by collagenase derived from Clostridium histolyticum.³⁶ However, dedicated in vivo studies are required to determine the feasibility, kinetics, and safety of this approach in clinical applications.

This proof-of-principle study suggests that scaffold-forming oligomeric collagen may address limitations associated with existing dermal fillers. The material’s tissue-matched mechanics, predictable localization, and favorable host response support further investigation of this platform for soft tissue volumization and restoration. Future studies will focus on extended longitudinal assessments, evaluation of in vivo reversibility (including controlled enzymatic degradation), injection volume-to-scaffold-volume predictability, and validation in clinically relevant anatomical models. Collectively, this work supports continued preclinical and translational development of scaffold-forming oligomer injectables for dermal filler and broader soft tissue reconstruction applications.

Authors’ Contributions

R.A.M. and C.P. contributed equally to this work. Conceptualization: R.A.M., C.P., and S.L.V.-H. Methodology: R.A.M., C.P., and S.L.V.-H. Investigation: R.A.M., C.P., L.Z., B.J., A.A., A.C., T.W., and P.F. Data curation: R.A.M., C.P., L.Z., B.J., A.A., A.C., T.W., P.F., and T.J.P. Formal analysis: R.A.M., C.P., L.Z., B.J., A.A., A.C., T.W., P.F., and T.J.P. Writing—original draft: R.A.M., C.P., and S.L.V.-H. Writing—review and editing: R.A.M., C.P., L.Z., B.J., A.A., A.C., T.W., P.F., T.J.P., W.G.C., S.L.H., and S.L.V.-H. Clinical insight: W.G.C. and S.L.H. Supervision: S.L.V.-H. Funding acquisition: S.L.V.-H. Project administration: S.L.V.-H. All authors approved the final article.

Footnotes

Acknowledgments

The authors thank the Weldon School of Biomedical Engineering Preclinical Studies Research Team for their assistance with the animal study. The authors also acknowledge the assistance of the Purdue University Histology Laboratory, a core facility of the National Institutes of Health-funded Indiana Clinical and Translational Sciences Institute.

Author Disclosure Statement

S.L.V.-H. is the founder and a shareholder of GeniPhys, Inc. and does not hold any current executive or operational role with the company. W.G.C. is a shareholder of GeniPhys, Inc.

Funding Information

This research was funded, in part, by generous donations provided by the McKinley Family Foundation.

Supplemental Material

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